2012 commissioning of a new total body irradiation protocol

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2012

Commissioning of a new total body irradiationprotocolZoe BaldwinUniversity of Wollongong

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Recommended CitationBaldwin, Zoe, Commissioning of a new total body irradiation protocol, Master of Science - Research thesis, Faculty of Engineering,University of Wollongong, 2012. http://ro.uow.edu.au/theses/3756

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COMMISSIONING OF A NEW

TOTAL BODY IRRADIATION

PROTOCOL

by

Zoë Baldwin

A thesis submitted in fulfilment of the requirements for the Master of

Science – Research degree

University of Wollongong

Faculty of Engineering

2012

ABSTRACT

COMMISSIONING OF A NEW TOTAL BODY IRRADIATION

TREATMENT PROTOCOL

by Zoë Baldwin

Centre for Medical Radiation Physics

The current planning process for adult Total Body Irradiation (TBI) using the

PLATO 2D treatment planning system (TPS) at the Royal Brisbane and

Women’s Hospital is cumbersome; it does not adequately simulate the

treatment technique and a basic manual calculation is used to account for the

effects of heterogeneities. A new treatment delivery technique is proposed;

this thesis addresses the commissioning of this new regimen, including the

acquisition of dosimetric data under treatment conditions to quantify the

accuracy achievable by performing dose calculations using Oncentra

MasterPlan TPS for this new treatment technique.

Two treatment planning calculation algorithms (a pencil beam and a collapsed

cone) and a total of five separate beam models (four of which were

formulated specifically for TBI) were examined. Comparisons were made

between calculated and measured data, which include percentage depth dose

(PDD) values, profiles, an output factor, doses within homogeneous water-

equivalent phantoms and doses within anthropomorphic phantoms.

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TABLE OF CONTENTS

Introduction ......................................................................................................................... 1 1.1 Total Body Irradiation (TBI) ....................................................................... 1 1.2 Treatment planning ....................................................................................... 2 1.3 Calculation algorithms .................................................................................. 5 1.3.1 Pencil Beam (PB) ........................................................................................... 5 1.3.2 Collapsed Cone (CC)..................................................................................... 6 1.4 Dosimetric verification with phantoms..................................................... 7 1.5 Considerations in the design of a TBI programme................................. 9 1.6 Complications and Organs At Risk (OAR) ............................................10 1.7 Compensation ..............................................................................................11 1.8 Technique and patient set-up ....................................................................14 1.8.1 Patient set-up ................................................................................................16 1.8.2 Complex patient set-up...............................................................................19 1.9 Beam energy..................................................................................................20 1.9.1 Dose build-up...............................................................................................21 1.10 Treatment Planning System (TPS) data...................................................21 1.10.1 Commissioning of TBI TPS Calculations........................................23

Confirmation of beam remodelling and data requirments PHASE I .....................25 2 Initial investigation..............................................................................................25 2.1 Patient positioning.......................................................................................25 2.2 Dosimetric verification ...............................................................................26 2.3 Small cubic phantom...................................................................................27 2.4 PDD acquisition ..........................................................................................29 2.4.1 PDD for MLC-delineated Collimator 45° beam...................................31 2.4.2 PDD comparison for MLC-delineated Collimator 45° beam.............32 2.4.3 PDD acquisition for Collimator 0°, 40 cm x 25 cm field ....................34 2.4.4 PDD comparison for Collimator 0°, 40 cm x 25 cm field ..................35 2.4.4.1 Absence of build-up screen ................................................................35 2.4.4.2 Bolus density settings...........................................................................37 2.5 Acquisition of beam profile free in air.....................................................39 2.6 Comparison of in air-profiles ....................................................................40 2.7 Cable signal ...................................................................................................42 2.8 Simple solid water phantom ......................................................................42 2.9 Simple solid water phantom comparison................................................43 2.10 Thorax phantom data acquisition.............................................................45 2.11 Thorax phantom comparison....................................................................47 2.12 Anthropomorphic phantom data acquisition.........................................48 2.13 Anthropomorphic phantom comparison................................................52 2.14 Outcomes......................................................................................................55 2.14.1 Nucletron data submission .................................................................55

Data model validation PHASE II ..................................................................................57

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3 Dosimetric verification ......................................................................................57 3.1 Output Factor...............................................................................................57 3.2 PDD...............................................................................................................59 3.2.1 Model 1 ..........................................................................................................62 3.2.2 Model 2 ..........................................................................................................64 3.2.3 Model 3 ..........................................................................................................65 3.2.4 Model 4 ..........................................................................................................67 3.3 Profiles ...........................................................................................................69 3.3.1 Model 1 ..........................................................................................................70 3.3.2 Model 2 ..........................................................................................................71 3.3.3 Model 3 ..........................................................................................................72 3.3.4 Model 4 ..........................................................................................................72 3.4 Simple phantom comparison.....................................................................73 3.5 Measurement consistency investigation ..................................................74 3.6 Thorax Phantom Comparison ..................................................................76 3.7 Anthropomorphic Phantom......................................................................78 3.7.1 TLD variation investigation.......................................................................80 3.7.1.1 TLD calibration at ESSD....................................................................81

Discussion and Conclusion.............................................................................................86 4 Project Overview ................................................................................................86 4.1 Discussion .....................................................................................................87 4.1.1 PDDs .............................................................................................................87 4.1.2 Beam profiles measured in air ...................................................................87 4.1.3 Output factor................................................................................................87 4.1.4 Thorax phantom..........................................................................................88 4.1.5 Simple phantom...........................................................................................88 4.1.6 Anthropomorphic phantom......................................................................88 4.1.7 General Discussion......................................................................................89 4.2 Conclusion and Future Work....................................................................90

Appendix.............................................................................................................................92 5 Appendices...........................................................................................................92 5.1 Phase 1 Pencil Beam TLD Results...........................................................92 5.2 Phase 1 Collapsed Cone TLD Results.....................................................93 5.3 Anthropomorphic phantom TLD comparison .....................................95 5.4 Anthropomorphic phantom TLD (results corrected for revised TLD calibrated in large phantom) comparison....................................................97

References...........................................................................................................................99

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ACKNOWLEDGMENTS

I would like to acknowledge my supervisors, Prof Anatoly Rozenfeld, Mr

Robert Fitchew, Dr Philip Back and Mr Darren Cassidy, for their generous

support and guidance, without which this project would not have been

possible.

I would like to thank the radiation oncology physics team and the total body

irradiation radiation therapist working group at RBWH CCS (past and

present) for their assistance and patience during the completion of this

project.

I would like to thank my parents for always supporting me in every aspect of

my life.

Finally I would like to thank my partner, for his dedicated patience,

understanding and un-wavering support.

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ABBREVIATIONS

AAPM American Association of Physicists in Medicine

ABUR After Build-Up Region

ACPSEM Australasian College of Physical Scientists and Engineers in Medicine

BMT Bone Marrow Transplant

BUS Build Up Screen

CAX Central Axis (of beam)

CCS Cancer Care Services

CCE Collapsed Cone Enhanced (calculation algorithm)

CT Computed Tomography

DVH Dose Volume Histogram

ESSD Extended Source to Surface Distance

IBUR In Build-Up Region

ICRU International Commission on Radiation Units and Measurements

IMRT Intensity Modulated Radiation Therapy

MU Monitor Units

OAR Organ At Risk

OMP Oncentra MasterPlan

PBE Pencil Beam Enhanced (calculation algorithm)

PDD Percentage Depth Dose

PTV Planning Target Volume

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RBWH Royal Brisbane and Women’s Hospital

RT Radiation Therapy

SSD Source to Surface Distance

SCD Source to Chamber Distance

TBI Total Body Irradiation

TLD Thermoluminescent Dosimeter

TPS Treatment Planning System

uA Type A standard uncertainty

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LIST OF FIGURES

Figure 1: An image of a digitised axial CT slice from the Plato 2D TPS, in which the lungs and spine have been outlined, however the true relative electron density (or physical density) from the CT data are not used in the Plato 2D dose calculation; default nominal values are used instead. ............................................................................................................................................3

Figure 2: The simulation position for the current treatment planning protocol, which will be used as the simulation and treatment position for the proposed new treatment technique. ..........................................................................................4

Figure 3: Lateral view of the current treatment set-up at the RBWH. The yellow lines indicate the 6 slice positions for the current treatment technique. .......................................................................................................................................4

Figure 4: An axial view of a treatment plan as displayed in Plato 2D. Note the bulk density corrections have been applied. ..........................................................................11

Figure 5: A jig used to hold the lead sheets employed in head shielding in TBI. .........................12 Figure 6: A simulation film taken of a patient for the lung compensator

design; this image is compared with a film taken at each treatment fraction for confirmation of treatment position. ..................................................................13

Figure 7: Chest plane from TBI plan, indicating the position of the bolus and its use in the treatment plan at RBWH...................................................................................14

Figure 8: Bolus used in a TBI treatment for both contour compensation and as a patient positioning aid. .......................................................................................................14

Figure 9: Lateral view of the current semi-reclining treatment position.........................................16 Figure 10: The proposed new supine treatment position..................................................................18 Figure 11: Demonstration of how the new technique can be utilised to treat

tall adult patients (pictured is a 194 cm “patient”), the treatment field is indicated by the white line and the dimensions of the BUS indicated by the green line. .........................................................................................................................19

Figure 12: The simple small cubic solid water phantom consisting of Gammex RMI 457 solid water slabs. ......................................................................................27

Figure 13: Dose plotted as a function of depth (corrected for effective point of measurement) in a 20 cm cubic solid water phantom, with and without the BUS..........................................................................................................................29

Figure 14: The PTW MP3 water tank used in the acquisition of the PDD data.................................................................................................................................................30

Figure 15: The phantom arrangement used to acquire the build-up region measurements for the PDD......................................................................................................31

Figure 16: The PDDs acquired at the patient plane in a large water tank at 380 cm SSD using the field delineated by the MLCs with a collimator angle of 45º. The plot labelled CAX was acquired 7 cm directly below the beam central axis (CAX); the other plot was acquired 43 cm away from the CAX plot in the gun-target direction. ....................................................................32

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Figure 17: Treatment plan configuration for the case of the PDD taken 7 cm directly below the beam axis in a 40 cm x 25 cm field.........................................................33

Figure 18: Treatment plan configuration for the case of the PDD taken 7 cm below the beam axis and 43 cm towards the target in a 40 cm x 25 cm field. ...............................................................................................................................................33

Figure 19: The PDDs acquired at the patient plane in a large water tank at 380 cm SSD using the field delineated by the jaws only with a collimator angle of 0º. The plot labelled CAX was acquired 7 cm directly below the beam central axis (CAX); the other plot was acquired 43 cm away from the CAX plot in the gun-target direction. ....................................................................34

Figure 20: The comparison of the PDDs measured 7 cm below the central axis and calculated by the treatment planning system in the absence of the BUS using both the pencil beam and collapsed cone algorithms, normalised to 10 cm depth. ......................................................................................................35

Figure 21: The comparison of the PDDs measured 43 cm away from the central axis and calculated by the treatment planning system in the absence of the BUS using both the pencil beam and collapsed cone algorithms, normalised to 10 cm depth. .................................................................................36

Figure 22: The comparison of the PDDs measured on the central axis and those produced by the treatment planning system using the collapsed cone algorithm, normalised to a depth of 10 cm, with three different density values applied. ................................................................................................................38

Figure 23: The comparison of the PDDs measured on the central axis and those produced by the treatment planning system using the pencil beam algorithm, normalised to a depth of 10 cm, with three different density values applied.................................................................................................................38

Figure 24: The y-axis profile obtained in air with a build-up cap with 50% of the central dose occurring at ±80 cm. ....................................................................................40

Figure 25: The y-axis profile normalised at the central axis position, calculated by the Collapsed Cone and Pencil Beam algorithms as compared with that measured...............................................................................................................................41

Figure 26: The simple solid water phantom replicated in the TPS for the 14 cm separation arrangement. ......................................................................................................44

Figure 27: Transverse view of the thorax phantom as set-up on the treatment bed. The mid-plane of the phantom is positioned 400 cm from the source (as indicated by the laser). The chamber is in the chamber cavity in the lung on the beam entrance side.........................................................................46

Figure 28: A beam’s eye view of the thorax phantom behind the BUS on the treatment couch...........................................................................................................................46

Figure 29: The treatment plan produced by OMP from a CT scan of the thorax phantom...........................................................................................................................47

Figure 30: The anthropomorphic phantom arrangement used in this investigation on the treatment bed. Lasers are used to align the phantom in the same manner as for a patient undergoing treatment...............................49

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Figures 31. a), b) c): Photos taken of a subset of the sections within the anthropomorphic phantom in which TLDs were placed to acquire dosimetric data.............................................................................................................................50

Figure 32.1 – 32.10: CT slices of the anthropomorphic phantom, showing the individual TLD positions within the phantom. The beam direction as it appears in the above figures is from the right. The percentage difference between doses at the TLD positions calculated by the collapsed cone and pencil beam algorithms and the corresponding measured doses are shown in blue and red respectively......................................................54

Figure 33: The water phantom used to acquire the output factor. Solid water, Perspex and plastic water are used to provide lateral scatter..............................................56

Figure 34: PDD as calculated at 7 cm below the CAX by the pencil beam algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose, in a large volume water phantom..............................60

Figure 35: PDD as calculated at 7 cm below the CAX by the collapsed cone algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose, in a large volume water phantom..............................60

Figure 36: PDD as calculated at 7 cm below and 43 cm lateral to the CAX by the pencil beam algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose ............................................................61

Figure 37: PDD as calculated at 7 cm below and 43 cm lateral to the CAX by the collapsed cone algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose ............................................................62

Figure 38: Treatment beam model PDD as calculated by the pencil beam algorithm for a 10 cm x 10 cm field normalised at dmax rather than the 40 cm x 25 cm field, showing better agreement with the measured PDD for the 40 cm x 25 cm field for models 2 and 3.........................................................68

Figure 39: Treatment beam model PDD as calculated by the collapsed cone algorithm for a 10 cm x 10 cm field normalised at dmax rather than the 40 cm x 25 cm field, showing better agreement with the measured PDD for the 40 cm x 25 m field for models 2 and 3...........................................................69

Figure 40: Profiles as calculated at 7 cm below the CAX by the pencil beam algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the central position ...........................................................................................69

Figure 41: Profiles as calculated at 7 cm below the CAX by the collapsed cone algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the central position. ..........................................................................................70

Figure 42: PDD as calculated by the pencil beam algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the depth of 10 cm...................................................................................................................................................77

Figure 43: PDD as calculated by the collapsed cone algorithm for each treatment beam model for a 40 cm x 25 cm field normalised to the depth of 10 cm. ...........................................................................................................................78

Figure 44: The graph displaying the average, minimum and maximum percentage deviations of the TPS calculated doses from the TLD doses for Pencil Beam calculations..........................................................................................79

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Figure 45: The graph displaying the average percentage deviations of the TPS calculated doses from the TLD doses for Collapsed Cone calculations. ..................................................................................................................................80

Figure 46: The anthropomorphic/ solid water phantom arrangement used in the TLD accuracy and linearity investigation. .......................................................................81

Figures 47.1 – 47.10: CT slices of the anthropomorphic phantom, showing the individual TLD positions within the phantom. The percentage differences from the TLD of doses calculated by collapsed cone and pencil beam algorithms are shown in blue and red respectively for model 0. The TLDs were calibrated in a large phantom at ESSD. ...................................84

Figure 48: The graph displaying the average, minimum and maximum percentage deviations of the TPS calculated doses from the recalibrated TLD doses for PBE calculations. ......................................................................85

Figure 49: The graph displaying the average, minimum and maximum percentage deviations of the TPS calculated doses from the recalibrated TLD doses for CCE calculations.......................................................................85

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LIST OF TABLES

Table 1: The gamma analysis results comparing the measured PDD and the Pencil Beam algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm)...............................................35

Table 2: The gamma analysis results comparing the measured PDD and the Collapsed Cone algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm)...............................................36

Table 3: The gamma analysis results comparing the 43 cm off axis measured PDD and Pencil Beam algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm).......................................37

Table 4: The gamma analysis results comparing the 43 cm off axis measured PDD and the Collapsed Cone algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build- up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm)..................................37

Table 5: The results of gamma analysis performed to compare the Pencil Beam calculation of the profile and the measured data.......................................................41

Table 6: The results of gamma analysis performed to compare the Collapsed Cone calculation of the profile and the measured data .......................................................42

Table 7: The doses as measured by converting the charge collected in the simple solid water phantom to dose, where the standard uncertainty (uA) is derived according to TRS398[72] as the standard deviation of the mean. ......................................................................................................................................43

Table 8: The comparison of doses calculated by OMP using the Model 0 data with the doses measured by the Farmer chamber in the simple phantom. The tabulated values are the percentages by which the calculated dose exceeds the measured dose. ..........................................................................44

Table 9: The doses to water in cGy derived from the charge collected from the chamber in the central tissue plug, the spinal bone plug and the central lung plug on the entrance side of the beam. Each measurement was corrected for extra signal from cable irradiation..................................47

Table 10: The percentage difference between doses measured by the ionisation chamber in the tissue, lung and bone cavities and those calculated by the treatment planning system for 500 MU. A negative difference indicates a lower calculated dose than the measured dose...............................48

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Table 11: The average doses measured by the TLDs at each point of interest in the anthropomorphic phantom, including the mean average dose and the mean standard deviation of doses measured...........................................................51

Table 12: The dose at the reference point in cGy/MU calculated by the treatment planning system for each beam model and calculation algorithm for a 40 cm x 25 cm field defined at the linac isocentre. ..................................58

Table 13: The dose at the calibration point in cGy/MU calculated by the treatment planning system for each beam model and calculation algorithm for a 10 cm x 10 cm field defined at the linac isocentre. The measured calibration value was 0.06661 cGy/MU for a 40 cm x 25 cm field. ........................................................................................................................................58

Table 14: Percentage difference between dose at the calculation point in cGy/MU calculated by the treatment planning system for each beam model and calculation algorithm and the measured value of 0.06661 cGy/MU at the calibration point. A negative difference indicates a lower calculated dose than the measured dose......................................................................59

Table 15: The gamma analysis results comparing the measured PDD below the central axis and the PBE Model 1 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm) ...............................................................................................................................................62

Table 16: The gamma analysis results comparing the measured PDD below the central axis and the CCE Model 1 PDD with 10% or 2% and 1 mm distance to agreement, regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm).......................................63

Table 17: The gamma analysis results comparing the measured PDD 43 cm off axis and the PBE Model 1 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm).......................................63

Table 18: The gamma analysis results comparing the measured PDD 43 cm off axis and the CCE Model 1 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm).......................................63

Table 19: The gamma analysis results comparing the measured PDD below the central axis and the PBE Model 2 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm). ..............................................................................................................................................64

Table 20: The gamma analysis results comparing the measured PDD below the central axis and the CCE Model 2 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-

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Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm). ..............................................................................................................................................64




Table 24: The gamma analysis results comparing the measured PDD below the central axis and the CCE Model 3 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1 mm to 500 mm). ..............................................................................................................................................66





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Table 29: The results of gamma analysis performed to compare the Pencil Beam calculation of the profile and the measured data for Model 1. ...............................70

Table 30: The results of gamma analysis performed to compare the Collapsed Cone calculation of the profile and the measured data for Model 1.................................71






Table 36: The percentage differences between the doses measured in the simple solid water phantom and those calculated using the PBE and CCE algorithms for all beam models investigated for the same phantom geometry and number of MU in OMP. The positive differences indicate higher calculated doses than the measured doses.............................73

Table 37: The percentage difference between doses measured by the ionisation chamber in the tissue, lung and bone and those calculated by the TPS for 500 MU, where the chamber cavity has been replicated in the TPS as being filled with water. The ionisation chamber readings have been corrected for cable signal. A negative difference indicates lower calculated dose than the measured dose......................................................................76

Table 38: The results of the comparison of TLD dose measurements in the anthropomorphic/ solid water phantom with doses measured by the ionisation chamber......................................................................................................................82

Table 39: The doses in cGy measured with TLDs within the anthropomorphic phantom and comparison with the existing beam model (model 0) doses calculated by the Pencil Beam algorithm in OMP. The “% diff TLD” is the percentage by which the calculated dose exceeds the measured dose..............................................................................................93

Table 40: The doses in cGy measured with TLDs within the anthropomorphic phantom and comparison with the existing beam model (model 0) doses calculated by the Collapsed Cone algorithm in OMP. The “% diff TLD” is the percentage by which the calculated dose exceeds the measured dose..............................................................................................94

Table 41: The doses calculated using the PBE algorithm for each beam model, at each TLD position, and a comparison with the doses measured by the TLDs. For each model, the first column shows the calculated dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the TLD-measured dose at that point for the same number of MU..................................................................................95

Table 42: The doses calculated using the CCE algorithm for each beam model, at each TLD position and a comparison with the doses

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measured by the TLDs. For each model, the first column shows the calculated dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the TLD-measured dose at that point for the same number of MU..................................................................................96

Table 43: The doses calculated using the pencil beam algorithm for all models and the percentage differences between the measured and calculated values. For each model, the first column shows the calculated dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the revised TLD-measured dose at that point for the same number of MU..........................................................................................97

Table 44: The doses calculated using the Collapsed Cone algorithm for all models and the percentage differences between the measured and calculated values. For each model, the first column shows the calculated dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the revised TLD-measured dose at that point for the same number of MU....................................................................98

C h a p t e r 1

INTRODUCTION

1.1 Total Body Irradiation (TBI)

Total Body Irradiation is a treatment technique that involves the irradiation of the whole

body. Several disorders where the entire body requires treatment have been proven to

respond to TBI, including lymphomas[1-3], leukemias[4-6] and aplastic anaemia[7]. In

situations where an allogeneic bone marrow transplant (BMT) or stem cell transplant is

prescribed it is common for such patients to undergo TBI in combination with chemical

therapy as part of a pre-transplant cytoreductive conditioning process[8, 9]. TBI yields two

beneficial outcomes in this process: firstly in the destruction of cancerous cells (also an

outcome of the chemotherapy) and secondly in the suppression of the immune system to

combat any rejection of transplanted bone marrow[10]. Although this preparatory

regimen for BMT has evolved from TBI alone, the combination with chemical therapy

agents has proven much more effective.

Advantages of TBI are that no organ is unintentionally spared and there are no sanctuary

sites, as the dose is homogeneous throughout and independent of blood supply[11]. The

radiation dose is not affected by any detoxification or excretion and if desired the dose

distribution may be adjusted through the use of partial compensators for the

radiosensitive normal tissues, and boosts may be prescribed for the radio-resistant regions.

As in all forms of patient therapy, in TBI it is desirable to achieve the maximum possible

therapeutic ratio[12] providing greater disease control and minimal normal tissue toxicity.

The radiation induced toxicity of normal tissue is a major limiting factor in the

effectiveness of total body irradiation. Clift et al[5] demonstrated the fine line between an

improvement in tumour control/ immuno-suppression and the transplant-related

complications for different irradiation regimens. A lower relapse rate for one regimen

translates to an improved long term survival, but with an increased risk of early mortality

resulting from increased acute toxicity.

2

1.2 Treatment planning

The use of Computed Tomography (CT)-based treatment planning systems in

conventional external beam radiation therapy is well established[13-16]. The tomographic

data provides the necessary information for the treatment planning system to determine

the patient’s external contour, to calculate the transmission and absorption of the

radiation in external beams, and hence to display the distribution of absorbed dose within

the patient. The CT data set which provides details of external contour, geometry of the

internal organs and their composition, constitutes a suitable base from which to calculate

the effect of tissue heterogeneities. Conventional external beam radiotherapy is conducted

at a nominal isocentric distance of 100 cm (often referred to as 100 cm Source-Axis-

Distance); treatment techniques are described as extended SSD when the source to

surface distance (SSD) is greater than 100 cm.

CT-based TBI treatment planning is less common. Because of the complex nature of an

extended SSD total body irradiation, Hui [17] investigated the accuracy of a CT-based

treatment planning system (Pinnacle™), finding that the use of CT data provides the

necessary anatomical and density information that ultimately results in a more accurate

dose calculation for TBI. This is especially important for the lung dose (an organ at risk),

where compensation is required to achieve the desired uniformity, and the CT data can

assist in compensator design. A Dose Volume Histogram (DVH) generated using the CT

data can provide homogeneity indexes for organs at risk (OAR) as well as the Planning

Target Volume (PTV).

Plato 2D (Nucletron, Veenendaal, The Netherlands) - the treatment planning system

currently used for planning TBI treatments at the Royal Brisbane and Women’s Hospital

(RBWH) - utilises 6 slices of CT data for the purpose of determining the location of the

kidneys and lungs (organs at risk) (see figure 3). The lung is outlined and a bulk density

correction (of 0.33 for adults and 0.41 for children) is applied. All other anatomical

structures are assigned a relative density of 1 (water). The external patient contours

obtained from the CT slices are used to account for the contour variation of the patient

(see figure 1). The technique is time consuming and does not reflect the dose to all

3

pertinent organs, nor the treatment position as the patient is simulated in a more supine

position (see figure 2) than the semi-reclining position used for treatment.

Figure 1: An image of a digitised axial CT slice from the Plato 2D TPS, in which the lungs and spine have

been outlined, however the true relative electron density (or physical density) from the CT data are not used

in the Plato 2D dose calculation; default nominal values are used instead.

The Plato 2D software is no longer supported by the manufacturer and therefore the use

of a new treatment planning system was proposed as the basis of a new TBI protocol.

Oncentra MasterPlan (OMP) (Nucletron, Veenendaal, The Netherlands) is also a CT-

based planning system that utilises helical CT data and is currently used for treatment

planning of all other three-dimensional conformal radiotherapy (3D-CRT) treatments at

the RBWH. In OMP, a 3D density matrix derived from pixel CT number values from CT

images or from user-specified density values assigned to regions of interest is generated,

and this is used by one of two well established[13, 18-22] calculation algorithms to

calculate doses to the patient,.

4

Figure 2: The simulation position for the current treatment planning protocol, which will be used as the

simulation and treatment position for the proposed new treatment technique.

Figure 3: Lateral view of the current treatment set-up at the RBWH. The yellow lines indicate the 6 slice

positions for the current treatment technique.

The prescribed doses range from 5 to 14 Gy delivered in up to 8 fractions, generally twice

daily. At RBWH doses for both paediatric and adult TBI are prescribed at the midpoint

5

behind the umbilicus. Dose limits for the critical organs are often included in the

prescription.

1.3 Calculation algorithms

Each commercial treatment planning system provides a selection of photon beam

treatment planning algorithms. Two algorithms available in Oncentra Masterplan for

photon beam treatment planning are Pencil Beam and Collapsed Cone. The same EGS4

Monte Carlo code proposed by Mackie et al [23] is used as the basis of the point dose

kernel for both algorithms. This point kernel is later used to obtain a pencil kernel (for the

pencil beam algorithm) through discrete integrations over the depth of calculations. The

accuracy and limitations of both calculation models have been investigated[19, 20, 22, 24-

32]. Oncentra Masterplan - unlike other commercially available treatment planning

systems - does not allow the user (physicist) to “tweak” the Monte Carlo generated energy

spectrum derived from the measured data. Instead this tweaking or beam characterisation

is conducted by the manufacturer off-site, based on the measured data supplied by the

user.

The dose calculation engine in OMP calculates the dose per incident energy fluence, based

on energy deposition kernels and combines this with calculated incident energy fluence

for the specified beam. The energy deposition kernel describes the dose response to an

incident radiation beam in water. For dose calculations the beams are treated separately

and the energy fluence exiting the treatment machine per beam is calculated as an initial

step of the dose calculation. The transport and dose deposition of the secondary particles

is taken into account by way of a pencil kernel for the pencil beam algorithm or a point

kernel for the collapsed cone algorithm.

1.3.1 Pencil Beam (PB)

This algorithm as implemented in OMP was originally proposed by Ahnesjö[32] and

utilises a particular technique to calculate the transport of photons and electrons and

deposition of dose. As the name suggests it is based on a pencil beam. For volumetric

integrations, the pencil beam is parameterised every 2.5 mm along the propagation

direction to a maximum of 502.5 mm. The actual beam is divided into a discrete grid of

6

infinitesimally narrow beams (pencils) which originate at the beam’s virtual source.

Equivalent path length corrections are applied to manage any regions of heterogeneity for

the primary dose contribution; however this does not take account of lateral transport of

secondary particles in heterogeneous media. A 1-dimensional convolution is applied along

fan lines with an exponential function to manage the scattered radiation[32-34]. The dose

at any point of interest is the result of the summation of the dose contributions from each

pencil beam. Heterogeneity corrections are performed by simply scaling the dose spread

function using the physical density in the path to save calculation time rather than

including density scaling for lateral heterogeneities. A decrease in the calculation accuracy

in regions where lateral electronic equilibrium is not present (where lung and air regions

are present) is the result of this simplification of the lateral transport of secondary particles

in the dose calculation. The penumbral widening observed in regions of low density (such

as lung) is not well predicted when using the pencil beam algorithm (the implementation

of which in OMP is known as Pencil Beam Enhanced PBE).

1.3.2 Collapsed Cone (CC)

Collapsed cone algorithms are gaining in popularity; as they provide a superior dose

calculation for regions where lateral charged particle equilibrium does not hold, such as in

lung and where the beam overshoots the patient’s body. The collapsed cone convolution

algorithm implemented in Oncentra Masterplan™ was first proposed by Ahnesjö[35]. It

involves apportioning all the energy released into coaxial cones of equal solid angles about

the point of interaction. The direct energy fluence and the head scatter energy fluence are

portioned into point kernels which describe the primary and scattered absorbed doses.

The dose at each point is obtained from interpolation in the 3D calculation grid. For

heterogeneities the radiological path length (as calculated from the density matrix) along

each cone axis replaces the geometrical path length. A limitation of the implementation of

the collapsed cone algorithm (known as Collapse Cone Enhance CCE) in OMP is in the

voxelisation of the calculation. In the case of the pencil beam algorithm the voxel

calculation size is determined by the user, however in the interest of calculation efficiency

the collapsed cone algorithm employs a finite number of voxels and as a result where the

volume is large (such as in the case of full body CT) the voxel size may be too large to

produce a dose distribution with high spatial resolution. This would limit the extent to

7

which small variations in the patient density or external contour would be handled by the

TPS calculation. A too coarse voxel size would be one that gives rise to significant errors

in reconstructing the positions of or dose distributions within OARs or targets. For TBI

voxel sizes of 5 mm or less should be satisfactory given the size of the OARs and the

PTV (the whole body).

1.4 Dosimetric verification with phantoms

The dose distribution of a clinical treatment may be inferred from measurements made in

a phantom. The term phantom refers to a material that is used to simulate the scattering

and absorption of radiation in tissue. Phantoms commonly consist of water, which has

similar absorption properties to muscle and soft tissue, however solid substances, such as

Plastic Water™[36] have been manufactured to also replicate the same properties.

Measurements of dose are made with dosimeters, the most commonly used type of which

in radiotherapy physics is the ionisation chamber, with various designs available[37-39]. A

common type of ionisation chamber (or ion chamber) consists of a conductive wall

surrounding a cylindrical gas-filled cavity within which is a centrally placed collecting

electrode. When a voltage is applied between this electrode and the wall, ions created in

the active volume by incident ionising radiation are attracted to the electrode and the

collected charge recorded by an electrometer. The primary measurement quantity is mass

ionisation (C/kg), the ionisation Q (C) per unit mass m (kg) of gas in the chamber, which

is related to dose (J/kg) to the gas in the chamber Dgas by the mean energy required to

produce one ion pair W (J). W/e (J/C) is then the mean energy expended per unit charge

produced, where e is the electronic charge (C), leading to the following equation:

=e

W

m

QDGas

The gas is most commonly air; more dense gases are sometimes used to increase

sensitivity by increasing the ionisation density within the volume[38]. One primary

requirement under which most radiotherapy dosimetry measurements are made is that in a

phantom the chamber is assumed to be small enough so as not to significantly perturb the

radiation field within the phantom. In such situations the charge produced and collected

8

in the chamber volume is proportional to the dose delivered (at the same point in the

absence of the chamber itself) to the phantom medium, from which the energetic

electrons producing the gas ionisation come in a megavoltage beam. Thus the dose to the

surrounding medium is given by the Bragg-Gray equation[39]:

GasMedGasMed SDD ,=

Where Smed,gas is the ratio of the mean mass stopping power of the energetic electrons in

the phantom medium and in the gas. For indirectly ionising radiation such as high energy

photons the contribution of any primary radiation to the ionisation events in the chamber

may be ignored. Since the theory assumes that all the measured ionisation is produced by

energetic electrons liberated by the photons in the surrounding phantom material, the wall

and other solid material of the chamber need to be similar to the phantom material and

the wall of chamber needs to be sufficiently thin, to ensure that almost all of the electrons

producing the charge measured have effectively been generated in the phantom.

Deviations from strict Bragg-Gray cavity conditions are dealt with by the use of

perturbation factors. In practice ion chambers are calibrated to provide chamber – specific

conversion coefficients which enable a direct conversion from measured charge to dose to

water.

Thermoluminescent Dosimeters (TLDs) have been used for in-vivo dosimetry[12, 14, 40].

TLDs provide a means of acquiring quick in-vivo measurements for radiotherapy and are

commonly used to verify the dose delivered to the eye when treating the head. Many

radiotherapy centres utilise TLDs for in-vivo measurement of superficial doses in total

body irradiation cases. There is a variety of TLDs available today, the most common

compositions being lithium fluoride, lithium borate and calcium fluoride, of which lithium

fluoride is the most widely used[41]. The most commonly used TLD in a radiotherapy

department is the Harshaw Chemical Company LiF TLD-100 made in the form of a small

disk. A TLD’s dose response may be described as the thermoluminescent output per unit

of absorbed dose in the phosphor and under certain conditions the dose response curve

of the LiF (TLD- 100) is linear up to 103 cGy and supralinear beyond this point[12]. The

TLDs can be calibrated so that their response is related to absorbed dose to water under

9

reference conditions (i.e. calibrated against an ion chamber). Metcalfe[12] outlines a cheap

and easy method for packaging TLD chips for clinical in-vivo measurements. Each centre

usually develops its own procedures for acquiring and processing TLD measurements

(which may include pre-irradiation annealing, post-irradiation annealing and a particular

heating program during the readout process). If these procedures are adhered to correctly,

they lead to accurate and consistent dose measurements and an uncertainty equal to 1

standard deviation of the mean value. The Australasian College of Physical Scientists and

Engineers in Medicine (ACPSEM)[15] recommends a calibration of the TLDs to be

undertaken with every measurement and a linearity check to be completed following the

initial use of each mode used or after a malfunction or repair of the TLD equipment.

Ionisation chambers and TLDs were used extensively in the work presented in this thesis.

1.5 Considerations in the design of a TBI programme

Historically TBI has been given as a single fraction, however the radiation induced side

effects can be better managed under a fractionated regimen, based on radiobiological

principles of preferential normal tissue repair[12]. Fractionation allows the prescription of

a higher total dose and improvement in the long term survival. Research by Deeg[42],

Cosset[43] and Cosset[44] has shown fractionated TBI can produce a better therapeutic

ratio than a non-fractionated scheme. According to Cox[45] the most common regimens

are 12 Gy given in six fractions over 6 days or twice a day over three days.

The aim of TBI is the ablation of bone marrow and the destruction of circulating

leukemic cells; since these are widely distributed throughout the body, the target for TBI

treatments is the entire body.

The achievement of uniform dose throughout the patient in TBI is a challenge; an

overdose to any critical organs will increase the toxicity while an underdose to the target

will increase the risk of relapse. A region of low density results in higher dose within and

beyond the low density region, as a result of the reduced attenuation of the beam as it

travels through the low density medium. The works of Keane et al[46] and Van Dyk et

al[47] have shown that a change of 5% in the dose to lung could result in a 20% change in

the incidence of pneumonitis, a complication potentially fatal for a whole lung irradiation.

10

Khan[36] discusses the clinical achievement of a high dose uniformity in TBI, accepting

dose uniformity along the body axis of ±10%, with extremities and certain non-critical

structures receiving higher doses, in agreement with AAPM report 17[8].

As survival rates have increased for patients treated with total body irradiation, the

investigation of long term side effects and the quality of life has become more significant

clinically. In TBI, treatment parameters prescribed such as the total dose, the fractionation

regimen[43, 44] and dose rate[48], are influenced by the possible toxicity to sensitive

organs and normal tissues such as lungs and kidneys. Consequently the accuracy attainable

in TBI treatment planning and dose delivery is a factor limiting the application of certain

techniques. The International Commission on Radiation Units and Measurements

(ICRU)[49, 50], recommends that the overall accuracy in the dose delivery of external

beam radiation therapy be ±5% and that the dose range within the target volume not

exceed -5% to +7% of the reference dose. However the American Association of

Physicists in Medicine (AAPM) Task Group 29[8] suggests that a 5% accuracy could be

unachievable in TBI, but if the prescribed dose falls far below the onset of normal tissue

toxicity levels then a 5% accuracy may be relaxed.

1.6 Complications and Organs At Risk (OAR)

In radiation therapy, to achieve the desired result, the destruction of cancerous cells needs

to occur without excessive destruction of normal tissues. Certain organs and tissues are

more radiosensitive than others and become a limitation when treating patients with

radiation therapy. It is important to limit the radiation dose to such organs so as to avoid

undesirable outcomes such as pulmonary toxicity.

In TBI the entire body is irradiated and as a result a number of critical structures can be

considered organs at risk. Limitations to the effectiveness of TBI in the conditioning

process exist, the most notable limitation being the normal tissue tolerance of certain

organs, such as the lungs, kidneys and eyes[8, 13, 14, 18]. Each of these is associated with

a higher tissue complication factor which results in a greater likelihood of cumulative

toxicity than other organs with lower complication factors[14]. In the prescribing and

planning of TBI, it is important to develop the appropriate procedures to protect these

11

organs at risk. The most concerning OAR is the lung, the associated complications for

which are one of the major causes of mortality following total body irradiation[18].

Figure 4: An axial view of a treatment plan as displayed in Plato 2D. Note the bulk density corrections have

been applied.

1.7 Compensation

Compensation is used for tissue density heterogeneities, surface irregularities and overall

separation variations. Because the large variations in body thickness and the presence of

tissue heterogeneities in TBI would lead to considerable non-uniformity of the dose along

the body, bolus and/or compensators are usually used to provide a much more uniform

dose distribution throughout the body.

Bolus material is often placed around the patient for support as well as contour

compensation. In regions where the patient thickness is significantly reduced, tissue

equivalent bolus may be added to achieve the required thickness. To account for

variations in patient shape either tissue-equivalent bolus or tissue compensators should be

employed. Although using the AP/PA treatment technique would reduce the amount of

12

the required lung tissue compensation, positioning of the arms to shadow the lungs in the

bilateral treatment technique partially compensates for lack of tissue homogeneity,

reducing the additional shielding and bolus needed for compensation[17, 51]. Bolus

materials generally employed in clinical use are pliable, soft tissue equivalent and non-

toxic. The equivalence refers to the volumetric electron density. A number of materials

have been used[8, 52]; bolus material if placed under the legs in TBI can provide both

added scatter and support during the treatment.

Figure 5: A jig used to hold the lead sheets employed in head shielding in TBI.

High density material such as lead is also utilised in the reduction of the intensity of

incident radiation to regions of OAR (see figure 5). For the regions of low electron

density such as the lung (a highly radiosensitive organ) the dose absorbed in the absence

of compensation would be higher than the intended dose, due to the lower attenuation of

the beam in the lung. Additional compensation to protect the lungs leads to an underdose

of the bone marrow in the sternum. At RBWH a megavoltage photon isocentric boost is

prescribed and delivered to the sternum at conventional SSD to address this issue.

13

Figure 6: A simulation film taken of a patient for the lung compensator design; this image is compared with a

film taken at each treatment fraction for confirmation of treatment position.

The current technique at the RBWH uses lead sheets (the thickness of which is manually

calculated based on measured transmission factors) to reduce the dose to the head and

lungs, positioned on a build-up screen (BUS) with their exact placement determined and

verified by film prior to each treatment. In addition bolus positioned adjacent to the

regions of the abdomen under the arms is used to provide additional attenuation to

reduce the dose to both the lungs and kidneys whilst also maintaining a relatively uniform

external surface. Although the bolus materials have densities a little different from water, a

bulk density of 1 g/cm2 is used when outlining the bolus material.

The majority of bolus material at used in TBI at RBWH is beeswax (see figure 8) with a

measured relative electron density of 0.96 (compared to water). Thicknesses of this bolus

used (up to 12 cm) are therefore unlikely to differ from water in their attenuation by more

than the attenuation produced by 0.5 cm of water. This equates to ~1.3% dose error

beyond the wax depth. For the current technique this could be an issue, however for the

proposed technique the patient will be CT scanned with the bolus material in place and

therefore taken into account in the TPS calculations.

14

Figure 7: Chest plane from TBI plan, indicating the position of the bolus and its use in the treatment plan at

RBWH.

Figure 8: Bolus used in a TBI treatment for both contour compensation and as a patient positioning aid.

1.8 Technique and patient set-up

There are a number of published techniques[9, 53-59] for the delivery of total body

irradiation. Each technique attempts to achieve uniform dose throughout the body while

not exceeding the tolerance dose to any OAR[14]. These techniques vary in the number

and type of the radiation sources used. Where most techniques use beams provided by

linear accelerators, use beams coming from multiple sources from several directions[36,

54]. A common and rather simple technique involves patients treated with opposing

bolus

15

fields at a fixed extended source to patient midline distance[36, 60], thereby employing

large, uniform photon beams to ensure the entire body receives a uniform dose.

Although it is possible to use multiple adjacent fields delivered at standard treatment

distances (100 cm SAD) to achieve total body irradiation, the use of such a technique

would result in complicated patient treatment set-ups and field junctions which introduce

non-uniformity and uncertainty in the dose distribution[36]. A further problem in using

this technique for TBI is that while the aim is to deliver radiation to the entire body, cells

circulating through the body may receive a reduced dose as a result of delivering smaller

fields separately. Thus an extended SSD treatment using opposed fields, which can easily

be implemented, is usually preferred if the treatment room is sufficiently large.

To achieve a field size large enough to encompass an adult human, the geometric

divergence of the beam from the source is utilised, which results in a required source to

patient mid-sagittal distance of 3 to 5 meters. Current conventional linear accelerators are

often used in large bunkers and as a result can apply a broad lateral horizontal beam (with

a gantry angle of 90º or 270º), with the maximum field size typically 40 cm x 40 cm (as

defined at 100 cm from the x-ray source), which diverges to 160 cm x 160 cm at 400 cm

from the source. The patient is set up at this distance with their mid-sagittal plane

perpendicular to the beam axis. The correct location of the mid-sagittal plane (patient

midline) is indicated by a ceiling-mounted laser. A collimator angle of 45º may be used to

take advantage of the longest available field dimension, which is 200 cm along the

diagonal at 400 cm from the source (the linac’s primary collimator limits the maximum

field diameter to 50 cm at 100 cm form the source). The opposing lateral field is achieved

by rotating the patient couch through 180º about a vertical axis. The technique currently

employed at RBWH utilises a conventional linear accelerator (Elekta Precise™) delivering

opposed lateral beams of 10 MV, with a BUS to avoid skin sparing.

16

Figure 9: Lateral view of the current semi-reclining treatment position.

Another advantage of performing the TBI treatment at an extended SSD is the

consequent reduction in dose rate to the patient which should be below 10 cGy/min to

prevent any dose rate related toxicity[48, 61]. This reduction occurs as a result of the

inverse square law. However if the dose rate is still not sufficiently low at the position of

the patient, a further reduction is required (achieved by decreasing the pulse repetition rate

of the linear accelerator). With the application of a decreased dose rate, an increase in the

treatment time results and the use of immobilisation devices may be necessary for certain

treatment protocols.

1.8.1 Patient set-up

In the case of extended SSD single source treatments, patients may be treated lying supine

with lateral beams; or standing[9, 62] or seated or lying on their side treated with an

anterior and a posterior beam[57]; in all instances it is necessary to rotate the patient in

order to obtain dose uniformity via an effective parallel opposed pair of beams.

In the cases where the maximum achievable field size is not sufficiently large, the patient’s

legs may be bent to reduce the overall length of the patient. Some techniques achieve this

by treating the patient seated in a custom chair; however such a treatment position may

17

not be replicated in the CT simulation, which can result in a reduction in the calculation

accuracy of the treatment planning process. In departments that have a sufficiently large

bunker it is common for patients to be treated lying supine on a bed. This has the

advantage of the patient being imaged for treatment planning and treated in the same

position. If the treatment requires the patient to have bent legs, with the use of a wide

bore CT the patient may still be imaged and treated in the same position, maintaining the

same pelvic tilt and anatomical geometry, potentially producing more accurate dose

estimates.

The Anterior-Posterior/ Posterior-Anterior (AP/ PA) technique requires appropriate lung

blocking to reduce the dose to the lung, concurrently allowing sufficient chest wall dose

coverage. Boost fields to regions of unacceptably low dose are often employed at the

RBWH as a part of the treatment process. In an opposed lateral supine treatment, the

patient’s arms are used as partial lung compensators. However with the AP/ PA

technique, if the patient is supported in an upright position, the overall separation of the

effective fields on the patient surface is decreased and less variable, which is advantageous

for dose uniformity, particularly for lower beam energies such as 60Co.

For reproducibility of set-up the supine position is superior[55]. However in using a

supine position in conjunction with a lateral beam (as is most commonly found) the dose

to the head would generally be too high, as a result of the greatly reduced lateral

separation. To reduce this high dose to the clinical volume a compensation filter (or

shield) is normally applied.

After consideration of the issues discussed above, the technique involving treatment of

supine patients with lateral beams, at a distance from the x-ray source to the patient’s mid-

sagittal plane of 400 cm, was proposed for the new TBI protocol at RBWH. However for

patients with a sufficiently large separation, consideration may be given to treating such

patients using an anterior-posterior technique to yield the required dose uniformity. In

order to maintain patients in such positions patient immobilisation and support systems as

complex as those used in the current treatment technique will be required.

18

The proposed technique (see figure 10) will enable the treatment of tall patients by

treating with the collimator rotated to 45º (see figure 11). In order to minimise dosimetry

complications resulting from scattered radiation from the direct irradiation of the steel

frame of the BUS and metal frame of the treatment bed, MLCs will be used to collimate

the treatment field to reduce the effect of this.

Figure 10: The proposed new supine treatment position.

19

Figure 11: Demonstration of how the new technique can be utilised to treat tall adult patients (pictured is a

194 cm “patient”), the treatment field is indicated by the white line and the dimensions of the BUS indicated

by the green line.

1.8.2 Complex patient set-up

Certain institutions have adopted more complex set-ups to address some of the issues

associated with TBI, to improve patient comfort and treatment reproducibility. Papiez[63]

utilised an extended SSD with the patient on the floor. The technique comprised 4 fields

in Anterior-Posterior (AP) and Posterior-Anterior (PA) treatments with the overlapped

regions modulated by multileaf collimators (MLCs).

Similarly to cater for a small bunker design, La Macchia[64], utilised an in-house designed

floor couch and segmented MLC fields, producing fields that were adjusted to deliver the

prescribed dose to the midline. An early technique utilised on 60Co units was the

translation of the couch during treatment, where the patient was moved continuously

through the treatment beam[59, 65].

20

Harden[66] describes a standing technique used at Addenbrooke’s Hospital, in which

patients are treated on a non-dedicated offset (relative to the centre of the bunker)

machine with anterior and posterior fields using large horizontal fields with customised

lung compensators.

1.9 Beam energy

Historically TBI was initially delivered with cobalt-60 low energy megavoltage (1.25 MeV)

beams[8] with protocol dose rates around 5 to 10 cGy/min. However 60Co machines have

largely been superseded by linear accelerators and total body irradiation techniques have

been adopted for linear accelerator beams*.

For lateral treatment techniques the photon beam energy required is dependent upon the

thickness of the patient and the dose homogeneity specification. The variation in

thickness of the patient along their sagittal axis will affect the homogeneity of dose

distribution; in addition the patient thickness along the central ray of the beam will also

impact upon the overall homogeneity. For the parallel lateral opposed pair technique, as

the patient thickness increases (or the beam energy decreases) the central axis maximum

dose near the surface increases relative to the midpoint dose, an effect referred to by

Khan[36] as the tissue lateral effect. In general for the sparing of subcutaneous tissues for

overall thickness exceeding 20 cm, Khan’s analysis indicates that 10 MV or higher energy

beams should be used.

The Task Group 29 of the AAPM[8] investigated the ratio of the peak dose to the

midline dose on the central ray versus the patient thickness, for 6- and 25-MV beams;

concluding that the greater the energy, the lower the dose variation throughout the

patient, the greater the treatment distance the lower the dose variation; the greater the

patient’s diameter the greater the dose variation. For a lateral opposed beam set-up where

the patients’ thicknesses range from approximately 38 cm to 50 cm only 25-MV x-rays at a

distance of 300 cm could result in a dose uniformity within 15% for the largest patients.

* in Australia there are no Cobalt- 60 units suitable for TBI

21

In summary 10 MV or higher energies are preferred, producing a more uniform dose

distribution for large separations in the case of the parallel lateral opposed pair technique.

They do however require a spoiler or bolus to increase the skin dose as discussed below.

The RBWH currently uses 10-MV beams and intends to continue the use of 10 MV (the

highest energy available on its current linacs) for future TBI treatments with the proposed

new protocol.

1.9.1 Dose build-up

It is well known that due to an initial electron fluence build-up with depth, the maximum

dose on the central axis of a megavoltage beam occurs not at the surface but at a certain

depth below it, with a progressive increase in dose from the surface to the depth of dose

maximum. The surface dose is typically about 20% of the maximum dose. Although this

skin sparing effect depends on a number of factors including SSD, field size and

configuration of any blocking tray present, in general the effect becomes more

pronounced as the photon energy increases. For use of higher energy x-rays, consideration

must be given to the effects of any low dose in the build-up region.

Because most TBI protocols do not require skin sparing, a “beam spoiler” is introduced

to bring the surface dose to at least 90% of the maximum dose; most of the electron

fluence build-up then takes place in the beam spoiler (or build-up screen), a sheet of low

atomic number material placed between the patient and the linac[8]. The thickness and

location of this BUS will be dependent upon the desired clinical dose criteria. In general

the screen is placed as close to the patient as possible, in order to achieve the greatest

increase in the skin dose over the whole surface of the patient facing the beam. Sanchez-

Nieto et al[67] state that a 1 cm thick PerspexTM sheet for a 10-MV photon beam will

increase the skin dose up to 97% of the maximum dose. A Perspex BUS of this thickness

is currently in use for TBI treatments at RBWH, and it is proposed to continue using it in

the new protocol.

1.10 Treatment Planning System (TPS) data

Treatment planning systems are used to calculate the dose to the patient, and need to do

this accurately if the desired outcome is to be achieved. They require data collected under

22

specified conditions to perform these calculations, and this data is normally collected

during the commissioning of the treatment unit and the TPS. Whether this data is

adequate for accurate calculations at an extended treatment distance is a matter of

conjecture.

The change in percentage depth dose (PDD) with increasing SSD is a consequence of the

inverse square law and the Mayneord Factor can be used to estimate this change[36].

However if this method is solely relied on for the dose calculations for patient dosimetry it

can lead to errors of up to 6% and therefore measurements of the PDD under treatment

conditions are recommended[36]. It should be noted for extended SSD TBI, Houdek et

al[68] compared the accuracy of dose calculations using beam data measured at the

extended SSD to that using data recalculated (from 100 cm SSD measured data) for a 10-

MV beam. Their limited investigation yielded comparable results for the two methods

which can be explained by considering the scattering conditions surrounding the patient

and water tank. A primary motivation or rationale for the investigation presented in this

thesis is the potential to improve upon the accuracy with which the current treatment

planning system calculates the dose distribution in TBI of a patient during treatment.

Phantom Scatter Factors (SP) can be applied to correct for the variation in the amount and

distribution of scatter resulting from a change in phantom size[36]. These factors quantify

the change in dose at a fixed point as the volume of phantom irradiated with a fixed

collimator opening is varied.

These factors used in dose calculation are normally interpolated from data acquired at 100

cm SSD, from measurements acquired using a phantom whose dimensions exceed the

field size; however for TBI this is not easily achievable and the TBI field is usually larger

than the phantom area presented to the beam. This difference in geometry would be

expected to impact any Tissue Maximum Ratios (TMRs) and field output factors for the

TBI calculations, however Curran et al[69] show that for TMRs measured in three

different phantom sizes at 100 cm SAD and 400 cm SAD there was at most a change of

1.5% in TMRs.

23

Collimator Scatter Factors (SC) are used to correct for the variation in the amount of

scatter in the beam incident on the phantom resulting from change in field size from the

calibration reference field[36]. Such factors are used in dose calculations and may be

derived empirically from interpolation of data acquired at 100 cm SSD. The variation

between SC measured at extended distance and 100 cm SSD should be limited[69] (for

field sizes larger than 25 cm x 25 cm of the order 0.3%[69]), however the SC measurement

at ESSD is potentially complicated by backscatter from adjacent surfaces and objects not

present in 100 cm SSD fields. This investigation aims to assess the suitability of utilising

the data available in the OMP measured at SSD 100 cm for accurate dose calculations for

400 cm source to mid-line distance.

In any case, measurements are needed at the extended distance to confirm the TPS

calculation accuracy under the treatment conditions.

1.10.1 Commissioning of TBI TPS Calculations

Following the selection of a treatment technique, the commissioning work begins.

According to Van Dyk[14] the beam data collected for conventional radiotherapy has little

application in the TBI setting; a separate set of dosimetric data is required for TBI

geometry. For a better prediction of dose by the treatment planning system, absolute

beam output calibration, percentage depth doses and beam profiles should be acquired, at

the extended SSD.

There are a number of dosimetric problems to consider in the commissioning of a TBI

protocol, specific to large field dosimetry which do not arise in conventional radiotherapy.

These issues are associated with the phantoms and ionisation chambers used in the

acquisition of the dosimetric data. Dosimetric inaccuracies can arise from the irradiation

of the ionisation chamber cable as a result of the relatively large radiation fields and

chamber leakage currents become more significant at the low dose rates typical of TBI.

Charge collection efficiency of the ionisation chamber will be greater at the low dose rate

and perturbation corrections used in dose calculations can depend on the source-chamber

distance (SCD)[14].

24

The selection of the phantom sizes for which this data is acquired is often considered to

be critical however for large distances Podgorsak[70] found that for output factors the size

of the phantom made a limited difference, although for the depth dose distribution there

was a significant variation in percentage doses at depths of 15 cm or greater. This effect

decreases slightly as the beam energy increases, attributed to the increase in the more

forward directed scattering of photons and a reduction in the amount of scatter produced

for higher energies.

25

C h a p t e r 2

CONFIRMATION OF BEAM REMODELLING AND DATA REQUIRMENTS PHASE I

2 Initial investigation

2.1 Patient positioning

The current treatment position involves the patient in a semi-reclining position proposed

by Khan[36] with knees elevated to reduce the overall length of the patient in the beam’s

eye view. The new position proposed in this thesis is designed to be more comfortable for

the patients during the lengthy treatment.

A survey of a number of RBWH TBI patients indicated a strong preference for the supine

position for comfort, confirming one of the critical benefits of the change in the

treatment protocol.

The overall height of the patient is an obvious limiting factor with regard to treatment

position. Although supine may be more comfortable, it makes it more difficult to fit a

very tall patient within the treatment field. In the event that a patient does not fit within

the size of the field, they will be treated with their knees bent. This position will be

maintained during both simulation and treatment. A low density patient support device (a

pillow) can be used to assist the patient in maintaining the desired position during the scan

and long treatment time, which will ultimately assist in the reproducibility of the treatment

setup. Because of the air gap between the BUS and the patient surface, the beam and the

screen need to extend some distance beyond the head and feet of the patient to produce

adequate build-up of electron fluence in these regions.

Although 40 cm x 40 cm is the largest field possible, the MLC system is used to reduce

the size of the field at the extended SSD to approximately the same dimensions as the

BUS. The collimator angle is set to 45º to take advantage of the maximum horizontal field

dimension of 50 cm (at 100 cm SSD) along a diagonal. The high density frame of the BUS

26

is kept outside the field, as OMP does not allow the simulation of materials with a density

exceeding 2.8 g/cm3.

2.2 Dosimetric verification

The first step in assessing the ability of a planning system to accurately model the

complexities of a treatment plan is to assess its ability to accurately model a simple case.

Dosimetric verification of a planning system can be achieved in many ways, the simplest

of which is the point measurement within a simple homogeneous phantom. Ionisation

chambers are used to collect charge deposited at a particular depth and the charge

collected can be converted to absorbed dose in a specified medium at that depth. The

charge collected is corrected for the influence quantities temperature and pressure,

polarity and recombination. An assessment was made to ascertain the need to modify the

parameters of the beam model, most notably the energy fluence kernels in OMP, to better

predict the doses at the extended distance: this was done by comparing a series of results

produced by OMP to experimental data, using the beam model already used clinically for

conventional (non-TBI) treatments in the TPS. This model will be referred to as Model 0

in this thesis.

Analysis performed using the Gamma Index as discussed by Low[71] is commonly

employed as a measure of the relevant difference between two dose profiles. For regions

of low dose gradients the difference between profiles is best assessed as the dose value

difference, in contrast to regions of high dose gradients (such as penumbra) where the

difference is best assessed by the distance-to-agreement for a particular dose value. As a

measurement of agreement, at each point along a profile the gamma index is given as a

combination of the dose and positional differences. When the gamma value exceeds 1 the

deviation between the two profiles is no longer within the tolerance limit set. The Gamma

analysis software employed in this thesis was developed in-house (using Matlab,

Mathworks, Natick, Ma, USA) based on the Low[71] evaluation method.

For the chambers used, the recombination effects are small (of the order of 1% of the

charge collected) at standard treatment distances and dose rates. The predominant

recombination effect component is approximately proportional to the dose (to the air in

27

the chamber) per pulse of radiation. At the extended (~400 cm) TBI treatment distance,

the recombination effects are negligible as the dose per pulse is lower by more than an

order of magnitude due to the inverse square law. The distance from the source to the

back wall is 495 cm.

2.3 Small cubic phantom

A small 20 cm cubic solid water phantom consisting of Gammex RMI 457 (Gammex

RMI, Madison, WI) slabs was used for a preliminary set of measurements due to the

geometrical simplicity and ease of set-up it provided. It was placed in the centre of the 40

x 40 cm MLC modified field at 389.75 cm SSD (400 cm to the centre of the phantom,

congruent with the clinical set-up).

The phantom comprised slabs of solid water, and a Farmer Type 2571 (Nuclear

Enterprises, Reading, England) ion chamber (used because of its compatibility with the

phantom – see Figure 12) connected to a PTW Unidos Webline (PTW-Frieburg, Frieburg,

Germany) electrometer was placed with its axis at 1 cm depth in the solid water. The slices

were re-arranged to allow the placement of the chamber at other depths while keeping the

SSD constant.

Figure 12: The simple small cubic solid water phantom consisting of Gammex RMI 457 solid water slabs.

To check the effect of the BUS (to be used during treatment), the charge was collected for

a series of depths with and without the BUS in the beam. The distance from the source to

the BUS is maintained at 380 cm.

28

The charge was converted to dose to water by applying a calibration factor. This factor

with units of cGy/nC was determined by acquiring the charge collected for a known dose

in the solid water phantom at dmax for a 10 cm x 10 cm field, at 100 cm SSD. This

process assumes that the spectrum of the large beam at the extended SSD is sufficiently

similar to that of the reference beam at 100 cm SSD that the mass stopping power ratio of

water or solid water to air is essentially the same for both measurements. The justification

for this assumption is as follows: (i) The effect of the differential attenuation of the beam

spectral components, by the additional 3 m of air is approximately equivalent to that of 3

mm of water, which would be negligible. (ii) The large area beam would have slightly

lower average energy than the 10 cm x 10 cm reference beam due to the increased

proportion of collimator scatter, but the effect of this would be no greater than for

conventional treatments at ~ 1 m SSD, where it is considered sufficiently small that no

correction is made for it in ion chamber dosimetry. (iii) The build-up with depth of

phantom scatter (Compton-scattered photons which have lower average energy than the

primary photons) reduces the average energy of a beam, and this reduction will increase as

the phantom becomes larger and generates more scatter. However the small cubic

phantom described in this section is of a size where again no correction is made in

conventional ion chamber dosimetry. (iv) The stopping power ratio of water to air varies

slowly with the variation in average energy in a 10-MV beam, so even a significant change

in average photon energy may produce an insignificant change in the response of the

chamber. The graphs in appendix 3 of the IAEA TRS-398[72] show that a 10% change in

the beam quality specifies TPR20,10, centred on the value 0.7 (the approximate value for a

10-MV beam), produces only a 1.5% change in Sw,air which is the predominant factor

affecting chamber response.

The results of these dose measurements within the solid water phantom are indicated in

figure 13. For all graphs presented in this thesis the error bars are insignificantly greater

than the thickness of the lines and are omitted for clarity.

29

10 MV, 389.75 SSD, 400 MUVarying depth in solid water at centre of 40 cm x 40 cm MLC modified field

0

5

10

15

20

25

30

0 2 4 6 8 10 12 14 16 18

Depth in solid water

Do

se (

cGy)

No screen

With screen

Figure 13: Dose plotted as a function of depth (corrected for effective point of measurement) in a 20 cm

cubic solid water phantom, with and without the BUS.

The results shown in figure 13 confirm the suitability of the BUS for increasing the

entrance dose and demonstrate that for a small volume phantom an average reduction in

local dose of 2.8% beyond the depth of maximum dose can be expected due to this BUS.

This is equivalent to shifting the PDD by ~ 1 cm towards the source for points beyond

the build-up region.

In a full patient-sized phantom the entrance dose would be a little larger than in this small

phantom due to increased backscatter. The build-up effect of the screen in the same sized

large field would be the same with a large phantom, so these preliminary measurements

indicated that the screen performed as anticipated and was likely to be satisfactory. This

was verified by later measurements in the build-up region of a large phantom with the

screen present - see section 2.4.

2.4 PDD acquisition

In order to acquire the pertinent data to generate a PDD, two phantom systems had to be

employed, one a large volume MP3 (PTW-Frieburg, Frieburg, Germany) computer

controlled water phantom system shown in figure 14 for the majority of the

30

measurements and the other a composite phantom consisting of Plastic Water and solid

water shown in figure 15 (CIRS (Computerized Imaging Reference Systems Inc.,

Norfolk, VA, USA) Plastic Water™ for build up and Gammex RMI 457 solid water for

backscatter) was used to acquire data for depths less than 52 mm. The second phantom

was required because of the physical limitations of the first phantom, for which the

minimum depth of measurement in a lateral beam was equivalent to 52 mm of water. The

attenuation coefficient of the Plastic Water when compared with real water in a 10-MV

beam was measured as 0.995 and applied in the calculation of water-equivalent depths in

this phantom.

Figure 14: The PTW MP3 water tank used in the acquisition of the PDD data.

The PDD data was acquired in both phantoms (regions of overlap between the two were

used to confirm no variation in response as a function of scatter and relative transmission)

using the same parallel-plate ionisation chamber PPC 40 (IBA Dosimetry,

Schwarzenbruck, Germany) so as to exclude any individual chamber response variations.

The effective measuring point of the ionisation chamber (1 mm below the front surface of

the chamber) was used for the PDD measurements.

31

Figure 15: The phantom arrangement used to acquire the build-up region measurements for the PDD.

The distance between the BUS and the phantom surface (17.8 cm) was maintained in this

PDD acquisition and the effect of varying this distance was not investigated.

The entrance dose which can be a significant parameter in TBI treatment techniques

would only decrease quite slowly by increasing this distance because of the very large field

size. Additionally in the treatment technique proposed the patient will be positioned on

the bed as close as practicable to the proximal edge of the bed and hence to the screen

which abuts the bed and will be less than the distance of 17.8 cm to the phantom. ICRU

[50] recommends that the dose within the target volume (in TBI this includes the skin)

should not vary more than +7% and -5% of the prescribed dose, so an acceptable value

for the entrance dose is considered to be at least 95% of the target dose. The entrance

dose measured in this phantom is greater than 96%.

2.4.1 PDD for MLC-delineated Collimator 45°°°° beam

Two separate PDDs were acquired, both at the extended SSD, one 7 cm below the central

axis of the beam and a second displaced from the first 43 cm parallel to the gun-target

direction (towards the target). A PDD is recommended by Andreo et al[73] as a part of

the commissioning process to verify the correct modelling of the energy spectra in a

treatment planning system.

32

PDD LA3, 400 MU, 7 cm below of CAX, 380 cm source-tank distance, 17.8 cm screen-tank separation

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 50 100 150 200 250 300 350 400 450 500 550

Depth (mm)

Per

cen

tage

Do

se

CAX

43 cm off axis

Figure 16: The PDDs acquired at the patient plane in a large water tank at 380 cm SSD using the field

delineated by the MLCs with a collimator angle of 45º. The plot labelled CAX was acquired 7 cm directly

below the beam central axis (CAX); the other plot was acquired 43 cm away from the CAX plot in the gun-

target direction.

As expected the beam is slightly less penetrating off axis as shown in figure 16. This

results from the lower initial energy and reduced beam hardening of the radiation further

off axis, (the latter due to reduction in the thickness of the flattening filter the off-axis

photons pass through), the obliquity of off-axis rays in the phantom and the decrease with

distance from the axis of the rate of build up of scatter with depth.

2.4.2 PDD comparison for MLC-delineated Collimator 45°°°° beam

In order to replicate the treatment geometry accurately an investigation was conducted to

assess the possibility of including the BUS as a region of interest to improve calculation

accuracy. An artificial CT image series was created to be sufficiently large to enable the

replication of the water tank within the image series in OMP. The BUS which has a

physical density of 1.19 g/cm3 had to be replicated in the CT data set as a bolus structure

(or region of interest).

Calculations were attempted however the treatment planning system was found to be

incapable of performing a calculation for a beam where the field size is set at 40 cm x 40

cm when the collimator is rotated away from 0º. This limitation of the planning system

33

was confirmed via correspondence with the manufacturer. Consequently any further

investigation of the large MLC-delineated collimator 45º beam was abandoned. Instead a

simple second treatment field was investigated, where the collimator is not rotated and the

field size is set to 40 cm x 25 cm.

Figure 17: Treatment plan configuration for the case of the PDD taken 7 cm directly below the beam axis in

a 40 cm x 25 cm field.

Figure 18: Treatment plan configuration for the case of the PDD taken 7 cm below the beam axis and 43 cm

towards the target in a 40 cm x 25 cm field.

34

2.4.3 PDD acquisition for Collimator 0°°°°, 40 cm x 25 cm field

Two separate PDDs were acquired at the extended SSD (380 cm to the tank) with the 40

cm x 25 cm field, the first 7 cm below the central axis and a second 43 cm from the first

at the same height.

PDD LA3, 400 MU, 7 cm below of CAX, 380 cm source-tank distance, 17.8 cm screen-tank separation

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 50 100 150 200 250 300 350 400 450 500 550

Depth (mm)

Per

cen

tage

Do

se

CAX

43 cm off axis

Figure 19: The PDDs acquired at the patient plane in a large water tank at 380 cm SSD using the field

delineated by the jaws only with a collimator angle of 0º. The plot labelled CAX was acquired 7 cm directly

below the beam central axis (CAX); the other plot was acquired 43 cm away from the CAX plot in the gun-

target direction.

The ratio of doses between the two curves (PDD on CAX and PDD 43 cm off CAX)

with the MLC delineated field and the jaw defined field differ by only 0.4 %. The

relationship between the CAX and 43 cm off axis PDD curves for the 40 cm x 25 cm

field is similar to that observed for the field delineated by the MLC system with the

collimator rotated to 45º.

35

2.4.4 PDD comparison for Collimator 0°°°°, 40 cm x 25 cm field

2.4.4.1 Absence of build-up screen

PDD Comparision, 400 MU, 7 cm below axis, 380 cm source-tank distance, 17.8 cm screen-tank separation

0%

20%

40%

60%

80%

100%

120%

140%

0 10 20 30 40 50 60

Depth (cm)

Per

cen

tage

Do

se

Measured PBE no build-up screen CCE no build-up screen

Figure 20: The comparison of the PDDs measured 7 cm below the central axis and calculated by the

treatment planning system in the absence of the BUS using both the pencil beam and collapsed cone

algorithms, normalised to 10 cm depth.

IBUR ABUR User defined

Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 28.73 4.47 143.67

G_mean 2.95 3.14 4.64

Pass rate (%) 52.81 0.04 1.36

Table 1: The gamma analysis results comparing the measured PDD and the Pencil Beam algorithm PDD

with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-Up Region

(IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region (range 1

mm to 500 mm).

The maximum gamma value (G_max) represents the worst agreement between the data

sets within the regions compared; the mean gamma value (G_mean) represents the mean

gamma value within the region compared. A value greater than 1 represents a failure to

fall within the limits set. The Pencil Beam model has an unacceptably low pass rate

(~53%) and will not be used for clinical use.

36


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 12.02 2.41 60.08

G_mean 2.20 1.52 1.64

Pass rate (%) 59.43 27.60 50.43

Table 2: The gamma analysis results comparing the measured PDD and the Collapsed Cone algorithm PDD

with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-Up Region


mm to 500 mm).

Although an improvement from the Pencil Beam model, the Collapsed Cone model too

has an unacceptably low pass rate (~28%) and will not be used for clinical use

PDD Comparision, 400 MU, 43 cm off axis, 380 cm source-tank distance, 17.8 cm screen-tank separation

0%

20%

40%

60%

80%

100%

120%

140%

0 10 20 30 40 50 60

Depth (cm)

Per

cen

tage

Do

se

Measured PBE no build-up screen CCE no build-up screen

Figure 21: The comparison of the PDDs measured 43 cm away from the central axis and calculated by the

treatment planning system in the absence of the BUS using both the pencil beam and collapsed cone

algorithms, normalised to 10 cm depth.

37


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 29.57 6.27 76.94

G_mean 3.01 4.22 5.89

Pass rate (%) 52.81 0.04 1.16

Table 3: The gamma analysis results comparing the 43 cm off axis measured PDD and Pencil Beam

algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-

Up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined

region (range 1 mm to 500 mm).


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 12.0164 2.4052 60.0822

G_mean 2.1984 1.5214 1.6428

Pass rate (%) 59.4306 27.6001 50.4299

Table 4: The gamma analysis results comparing the 43 cm off axis measured PDD and the Collapsed Cone

algorithm PDD with 10 % or 2 % and 1 mm distance to agreement. Regions of interest include the In Build-

up Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined


From the above comparisons it is apparent that the data in the TPS based on data

collected at 100 cm SSD will be insufficient to accurately predict the doses at the ESSD by

using either PBE or CCE.

2.4.4.2 Bolus density settings

It was hypothesised that the OMP, operating on a model derived from measurements

made on the 10-MV Elekta (Elekta, Stockholm, Sweden) Precise beam at conventional

treatment distances, would be sufficiently accurate for TBI provided a structure replicating

the BUS is included in the calculation. This structure was included in the TPS plan as a 1

cm thick bolus region of interest (ROI). Although the physical density of this screen was

known, an investigation was conducted in order to assess if this structure could be

38

tweaked to force the simulated beam PDD to match the measured data. This was

performed by altering the density of the screen.

Percentage Depth Dose 7 cm below CAX, Measured data and Collpased Cone data comparision

0%

20%

40%

60%

80%

100%

120%

140%

0 5 10 15 20 25 30 35 40 45 50

Depth (cm)

Per

cen

tage

Do

se

Measured

Bolus 1 g/cm3

Bolus 1.2 g/cm3

Bolus 1.5 g/cm3

Figure 22: The comparison of the PDDs measured on the central axis and those produced by the treatment

planning system using the collapsed cone algorithm, normalised to a depth of 10 cm, with three different

density values applied.

Percentage Depth Dose 7 cm below CAX, Measured data and Pencil Beam data comparision

0%

20%

40%

60%

80%

100%

120%

140%

0 5 10 15 20 25 30 35 40 45 50

Depth (cm)

Per

cen

tage

Do

se

Measured

Bolus 1 g/cm3

Bolus 1.2 g/cm3

Bolus 1.5 g/cm3

Figure 23: The comparison of the PDDs measured on the central axis and those produced by the treatment

planning system using the pencil beam algorithm, normalised to a depth of 10 cm, with three different

density values applied

39

From the results presented in figures 22 and 23 above the maximum deviation appears to

be ~6 % of dMax. at ≥ 35 cm depth where the PDD is ~ 40%. This would correspond to

%100*140

6~

which is equal to ~ 4% deviation for a bilateral treatment, which could be considered

clinically acceptable in TBI.

Although the inclusion of the ROI replicating the BUS in treatment plans significantly

improved agreement with measured data, it was deemed too time consuming and in

certain clinical cases would prove to be unachievable, as OMP restricts the calculation of

dose to within the CT data set; any structure outside of the CT data is ignored. For the

case of large patients the inclusion of such a ROI would not be possible. The aim of this

project is to enable accurate calculations for all TBI patients at RBWH, not just a subset.

For both of the investigations described in this section, the current 10-MV model does

not predict the dose variation with depth correctly within the tolerance values set for each

investigation respectively, leading to the conclusion that an attempt to refine the method

of dose calculation is warranted. It is expected that generally the dose calculation should

be accurate to within ± 5 % of the measured data.

2.5 Acquisition of beam profile free in air

In order to investigate the overall limitation of the TPS in the cross beam direction at a

depth beyond the build-up an in-air profile was measured. A Farmer type ion chamber

fitted with a small Perspex build-up cap (5 cm in diameter) to produce charged particle

equilibrium was clamped to a retort stand. Charge measurements were acquired at the

height of the treatment beam isocentre, at various points across the 40 cm x 25 cm field.

40

10 MV, 400 cm SCD, 40 cm x 25 cm field, Y- Profile

3

3.5

4

4.5

5

5.5

6

6.5

7

7.5

8

-110 -90 -70 -50 -30 -10 10 30 50 70 90 110

Position on Y axis, at height of the beam axis

Ch

arge

co

llect

ed (

nC

) as

a f

un

ctio

n o

f M

U

Figure 24: The y-axis profile obtained in air with a build-up cap with 50% of the central dose occurring at

±80 cm.

The above investigation was conducted to check the symmetry and flatness in air at the

extended distance. The resultant y-axis profile was also expected to be useful in

interpreting the measured and TPS–calculated doses at off-axis points covering the length

of the patient. In the x-axis direction only a very limited range of the beam area will be

used, covering the anterior-posterior dimension of the patient, therefore this data was not

collected separately.

2.6 Comparison of in air-profiles

The phantom geometry (the build-up cap) was replicated in the treatment planning

system; the profiles produced by each calculation algorithm were extracted.

41

Profile Collapsed Cone, Longitudinal profile400 cm Source to chamber distance, 40 x 25 cm field

25%

35%

45%

55%

65%

75%

85%

95%

105%

115%

-100 -80 -60 -40 -20 0 20 40 60 80 100

Position (cm)

Per

cen

tage

do

se

Measured CCE Model 0 PBE Model 0

Figure 25: The y-axis profile normalised at the central axis position, calculated by the Collapsed Cone and

Pencil Beam algorithms as compared with that measured.

Gamma analysis was performed to quantify the differences between the profiles calculated

by the TPS and the measured profile:

Penumbra Penumbra Inner

Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 1.47 1.28 1.25

G_mean 0.27 0.35 0.33

Pass rate (%) 96.53 94.19 98.81

Gamma norm type Locally Central axis Central axis

Table 5: The results of gamma analysis performed to compare the Pencil Beam calculation of the profile and

the measured data.

42

Penumbra Penumbra Inner

Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 0.65 1.61 8.41

G_mean 0.16 0.29 0.54

Pass rate (%) 100.00 98.84 92.68


Table 6: The results of gamma analysis performed to compare the Collapsed Cone calculation of the profile

and the measured data

The above comparison favours the collapsed cone calculation, however the beam profile

is not considered a major factor in selecting a suitable calculation beam model. The

differences are only significant near the edges of the field where the beam will in most

clinical cases overshoot the patient.

2.7 Cable signal

The inclusion of the ionisation chamber cable in the path of the beam can result in a

spurious component of the measurement which is known as cable signal. To quantify this

signal, the chamber was irradiated under reference conditions (100 cm SSD for a 10 cm x

10 cm field) with a standard length of cable within the path of the beam and the charge

recorded; this was repeated this time including an equivalent length of cable as would be

in the path of the beam for the measurements made at 400 cm SCD. The percentage

variation in signal was recorded as 0.8% for the approximate length of cable that would be

present in the treatment beam for the Farmer chamber in the simple solid water phantom

described in section 2.8. The cable effect constitutes a reduction of signal 0.8% for the

Farmer chamber and an increase in signal of 7% for the Pinpoint chamber (used in

following chapters).

2.8 Simple solid water phantom

A small solid water phantom similar to the phantom shown in figure 12 was also used in

the investigation of the variation of the midline dose as a function of separation (thickness

in the beam direction). This was performed to assess the system’s ability to correctly

predict the dose at the prescription point for a variety of separations. The phantom was

43

placed on the treatment bed at constant 400 cm source to chamber distance, with the

chamber reference point on the beam axis, and doses were measured at this point, while

the phantom thickness surrounding the chamber was increased.

The ion chamber readings were converted to dose for 400 MU delivered. This process

assumes that the chamber perturbation factors in solid water do not differ significantly

from those in water; which are accounted for by the kQ factors in the calibration chain†.

Depth (cm) Dose (cGy) Stdev uA

5 24.74 0.01 0.01

14 19.90 0.01 0.01

16 18.90 0.01 0.01

21 16.42 0.01 0.01

26 14.23 0.01 0.01

Table 7: The doses as measured by converting the charge collected in the simple solid water phantom to

dose, where the standard uncertainty (uA) is derived according to TRS398[72] as the standard deviation of

the mean.

The dose on the phantom side opposite to the beam incident surface (also known as the

exit dose) was looked at as part of a comprehensive dosimetry investigation (section 2.12).

2.9 Simple solid water phantom comparison

The simple solid water phantom geometry was replicated in the TPS, by creating an empty

image CT series and recreating the phantom as a region of interest with a defined density

of 1 g/cm3. The dose was calculated using both pencil beam (PBE) and collapsed cone

(CCE) calculation algorithms at the midpoint for 400 MU. This plan does not include the

BUS; as discussed previously the inclusion of this screen in the plan is not clinically

achievable.

† For more details see TRS-398[73]

44

Figure 26: The simple solid water phantom replicated in the TPS for the 14 cm separation arrangement.

The doses calculated by the TPS are compared with the doses measured in the solid water

phantom and the differences are found in table 8.

PBE CCE PBE – CCE

Depth Model 0 Model 0 Model 0

5 7.3% 4.8% 2.5%

14 8.6% 5.2% 2.4%

16 8.7% 5.8% 2.9%

21 9.1% 5.3% 3.8%

26 8.5% 6.6% 1.9%

Table 8: The comparison of doses calculated by OMP using the Model 0 data with the doses measured by

the Farmer chamber in the simple phantom. The tabulated values are the percentages by which the

calculated dose exceeds the measured dose.

From the above data although the percentage depth dose for this beam model is not

considered clinically acceptable for dose calculations, when calculating the dose for this

phantom the collapsed cone algorithm is sufficiently accurate; 2.8% of these differences

can be attributed to the absence of the screen in the calculations.

45

2.10 Thorax phantom data acquisition

A thorax phantom (Computerized Imaging Reference Systems Inc., Norfolk, VA, USA)

was used to verify the dose (to water) in tissue equivalent, bone equivalent and lung

equivalent materials, using a small volume ion chamber‡. This chamber was used because

the phantom chamber cavity had been milled for this particular chamber; a chamber with

a larger volume would have been preferred to improve the signal to noise ratio and reduce

the effect of the cable signal.

The thorax phantom was arranged on the treatment bed as pictured in Figures 27 and 28,

with the mid-sagittal plane of the phantom positioned 400 cm from the source (30 cm

from the surface of the BUS to the mid-sagittal plane) and the highest point on the

anterior surface of the phantom several centimetres below the beam central axis. The

distance from the wall to the closest side of the phantom was ~ 82 cm. The 40 cm x 25

cm beam was used to deliver 500 MU. The ion chamber was placed in cavities milled

within plugs which fit in the phantom at key points (as shown in figure 27). These plugs

are made from substances that replicate the physical properties of lung, bone and soft

tissue. Plugs for this phantom that do not contain chamber cavities are used where the

chamber is not required and are also employed in the simulation process, that is, when a

CT is obtained of the phantom the chamber is not present and so does not perturb the

subsequent dose calculation.

‡ PTW Pinpoint (PTW-Frieburg, Germany) 31016 chamber, nominal sensitive volume of 0.016 cm2

46

Figure 27: Transverse view of the thorax phantom as set-up on the treatment bed. The mid-plane of the

phantom is positioned 400 cm from the source (as indicated by the laser). The chamber is in the chamber

cavity in the lung on the beam entrance side.

Figure 28: A beam’s eye view of the thorax phantom behind the BUS on the treatment couch.

47

Central tissue Lung entrance side Central bone

Average 30.09 34.67 30.29

Stdev 0.01 0.04 0.01

uA 0.01 0.02 0.01

Table 9: The doses to water in cGy derived from the charge collected from the chamber in the central tissue

plug, the spinal bone plug and the central lung plug on the entrance side of the beam. Each measurement

was corrected for extra signal from cable irradiation.

The ion chamber readings were converted from charge collected, to dose to water using a

conversion factor with units of cGy/nC obtained by acquiring the charge collected under

reference conditions for a known delivered dose. In all cases the values measured were

corrected for extra cable signal.

2.11 Thorax phantom comparison

The phantom was scanned, a plan was created replicating the experimental conditions

(without the BUS in place) and the doses were calculated using OMP to the regions where

the ion chambers were placed using 500 MU.

Figure 29: The treatment plan produced by OMP from a CT scan of the thorax phantom.

A comparison of the ionisation chamber measured doses with those calculated by the

treatment planning system for both the Pencil Beam and Collapsed Cone algorithms was

48

carried out, the results of which are presented in the table 10 . It should be noted that

doses quoted in table 10 are those obtained for dose to water within these bone, tissue

and lung substances. In order to calculate the doses to these materials directly a ratio of

the mass energy absorption coefficients for each material to water would need to be

applied, however the exact composition of this phantom was unknown and applying these

factors would introduce further uncertainty. Consequently, in the TPS the density of the

chamber volume (including the wall) was set to 1 g/cm2 (water).

PBE CCE

Lung Bone Tissue Lung Bone Tissue

-2.9% -9.1% -10.5% -5.7% -3.2% -4.3%

Table 10: The percentage difference between doses measured by the ionisation chamber in the tissue, lung

and bone cavities and those calculated by the treatment planning system for 500 MU. A negative difference

indicates a lower calculated dose than the measured dose.

From the tabulated dose deviations in table 10, the ratio of doses calculated within bone

to those within tissue are slightly higher when calculated by PBE than by CCE. This

indicates a difference in the energy kernel for the heterogeneity correction. The overall

magnitude of the deviations between the measured and calculated doses for both PBE

and CCE is not consistent with the previous investigation(s) presented in section 2.9; the

deviation is a calculated dose that is much lower than the measured dose.

2.12 Anthropomorphic phantom data acquisition

For a more complex investigation a RANDO (Alderson Research Laboratories Inc.,

Stanford, CT, USA) anthropomorphic phantom is used, which includes materials that

simulate muscle, bone and lung and has the contours of an adult male from the head to

the upper thighs. TLDs can be used to verify the accuracy of the TPS in calculating dose

in the presence of heterogeneities and contour variations. TLD-100 chips (Harshaw

Chemical Co., Solon, Ohio, USA) were used as a convenient dosimeter to be placed

within the anthropomorphic phantom. This phantom was purchased in the 1970s and

contains real human bone, (which however is quite porous from osteoporosis).

49

TLDs were placed at several key positions such as the organs at risk (two at each point to

improve the statistical uncertainty) within the anthropomorphic phantom shown in figure

30. Any gaps around the TLDs were filled with tissue-equivalent bolus. Figures 31.a, 31.b

and 31.c show a few of the sections in which TLDs were placed.

The phantom was placed on the treatment bed and aligned so that the mid-sagittal plane

was 400 cm from the source (and 30 cm from the BUS surface), and irradiated by a single

40 cm x 25 cm beam (incident on the right side of the phantom) for 1000 MU. This

procedure was repeated once to improve measurement statistics (for a total of 4 TLD

readings at each point of interest). A single-sided irradiation was used rather than a

bilateral delivery as this was considered a more sensitive way to evaluate the calculation

model and algorithms.

Figure 30: The anthropomorphic phantom arrangement used in this investigation on the treatment bed.

Lasers are used to align the phantom in the same manner as for a patient undergoing treatment.

50

a) b) c)

Figures 31. a), b) c): Photos taken of a subset of the sections within the anthropomorphic phantom in which

TLDs were placed to acquire dosimetric data.

The TLD measurements were converted to dose by applying a calibration factor acquired

under reference conditions and are tabulated in table 11.

The TLD measurement uncertainty at each point (type A[72]) is quantified by the

standard uncertainty (uA). Which is the standard deviation of the mean value of the four

dose measurements at the point.. These values, together with the average measured dose

taken over all points and the average value of uA, are shown in table 11.

51

Dose Point Dose (cGy) Stdev (uA)

1 56.09 2.23 1.115

2 59.68 0.92 0.46

3 66.92 3.71 1.855

4 53.28 0.1 0.05

5 65.58 2.42 1.21

6 68.29 1.32 0.66

7 56.01 0.95 0.475

8 34.29 0.16 0.08

9 32.47 0.5 0.25

10 44.24 0.54 0.27

11 51.28 0.72 0.36

12 56.37 1.85 0.925

13 74.69 0.63 0.315

14 72.92 0.77 0.385

15 40.02 0.41 0.205

16 47.07 1.49 0.745

17 57.72 0.47 0.235

18 65.03 2.01 1.005

19 75.18 1.38 0.69

20 40.08 0.97 0.485

21 35.12 0.09 0.045

22 61.56 1.65 0.825

23 44.25 1.17 0.585

24 53.64 1.37 0.685

25 49.2 0.37 0.185

26 40.44 0.77 0.385

27 59.82 0.54 0.27

28 45.06 0.74 0.37

29 44.44 1.28 0.64

30 46.64 2.35 1.175

31 64.93 2.55 1.275

32 66.04 4.03 2.015

33 35.49 0.28 0.14

34 43.49 1.08 0.54

35 53.32 1.35 0.675

36 64.1 2.18 1.09

37 73.29 0.27 0.135

38 59.4 1.06 0.53

39 57.28 2.19 1.095

40 60.35 0.36 0.18

41 44.48 0.11 0.055

average 54.13 1.2

Table 11: The average doses measured by the TLDs at each point of interest in the anthropomorphic

phantom, including the mean average dose and the mean standard deviation of doses measured.

52

2.13 Anthropomorphic phantom comparison

The anthropomorphic phantom was scanned at 2 mm slices and the CT images (2 mm

thick) were imported into the planning system and the doses calculated for a plan

arrangement of a single beam directed at the phantom with a mid-sagittal distance of 400

cm from the source, using the 40 cm x 25 cm field size and 1000 MU. This investigation

tests the overall uncertainty in the simulation, treatment planning and dose calculation

process. The reproducibility of set-up may be estimated from examination of the two sets

of TLD data acquired from the anthropomorphic phantom. The IAEA[73] recommend a

final method of verification of the overall dose calculation accuracy of a TPS, in which an

anthropomorphic phantom is simulated, doses calculated and doses measured under

treatment conditions.

For each point replicating the TLD an equivalent small volume structure was substituted

in the TPS with a density of 1 g/cm3, so that the TPS calculates the dose to water, which

is the quantity the TLDs are calibrated to indicate. The calibration of the TLDs is

performed at the isocentre by delivering a known dose under reference conditions to a

subset of a batch at the same time of the phantom irradiation. Individual scaling factors

are applied to each TLD to convert the thermoluminescence recorded in readout to dose.

The results of comparison of the dose calculations with the dose measurements are

tabulated in tables 39 and 40 in the appendix.

The following series of images are taken from the CT slices of the anthropomorphic

phantom and display percentage dose differences between the doses calculated by each

OMP algorithm (pencil beam represented in red and collapsed cone in blue) and the

corresponding measured doses.

54

Figure 32.1 – 32.10: CT slices of the anthropomorphic phantom, showing the individual TLD positions

within the phantom. The beam direction as it appears in the above figures is from the right. The percentage

difference between doses at the TLD positions calculated by the collapsed cone and pencil beam algorithms

and the corresponding measured doses are shown in blue and red respectively.

The results of the average of the TLD measurements at each position obtained in the

anthropomorphic phantom for a single beam irradiation, show that if the conventional

10-MV model (aka model 0), calculated using the Pencil Beam algorithm, were to be used

clinically, an underestimate of the dose of up to 13% and an overestimate of up to 15%

would occur. In the case of the Collapsed Cone algorithm an underestimate of up to 12%

and overestimate of up to 9% would occur. For all points the doses calculated by the

Collapsed Cone are lower than those calculated by the Pencil Beam algorithm. The

average deviation of the calculated doses from the measured doses was 7% for PBE and -

11.7% for CCE. The exit doses for both models are noted as having superior agreement

with measured data when the beam does not pass through regions of heterogeneity,

consistent with percentage differences observed in the PDD comparison (section 2.4.4).

55

In both cases the discrepancies were deemed too great for clinical implementation of TBI

treatments.

Although a single beam configuration (single beam irradiation) is investigated in the above

comparisons, the effect of rotating the phantom and performing a second irradiation

would increase the measurement set-up error. Using an opposing beam configuration

would reduce the discrepancies within the build-up region.

2.14 Outcomes

Investigations were conducted into the dose calculation accuracy of both algorithms of

the current treatment planning model in regions of heterogeneities, and as a function of

separation and as a function of depth in a homogeneous phantom. From these

comparisons it was concluded that the current treatment planning model as it stands does

not yield sufficiently accurate calculations of dose for the clinical implementation of

extended SSD treatments in OMP at RBWH. It was conjectured that with better

characterisation of the beam, improved agreement with measured data could be achieved

by the TPS, particularly if the effect of the BUS could be incorporated in the beam model,

which would have the added advantage of simplifying and increasing the scope of

treatment planning.

To enable them to improve the beam model for TBI calculations, the manufacturer

(Nucletron®) requested certain data which included a PDD of the treatment field to be

used at the extended SSD and an absolute dose acquired under the same treatment

conditions.

2.14.1 Nucletron data submission

The central axis PDD acquired as described in section 2.4.3, was submitted. For the

output factor, the same beam set-up as that used to acquire the PDD data (40 cm x 25 cm

at isocentre, 10 MV) was directed at a smaller volume water phantom. However, this

phantom was centred on the central axis of the beam. A NE 2571 Farmer ionisation

chamber was placed on the beam axis with its reference point at a depth of 10 cm in the

phantom. The charge collected for 400 MU delivered was converted to absorbed dose,

according to the IAEA TRS-398 dosimetry protocol formalism.

56

The phantom dimensions for the output measurements were approximately 60 cm x 30

cm normal to the beam direction x 40 cm in the beam direction. The phantom is shown

in figure 33. The phantom size was different from that of the large automated water

phantom used to obtain the PDD data.

Figure 33: The water phantom used to acquire the output factor. Solid water, Perspex and plastic water are

used to provide lateral scatter.

The phantom was placed upon the intended treatment bed, which consists of an ordinary

hospital bed and a thick hospital mattress. The metal frame of the bed was not

encompassed within the light projected field size, so as to minimise any effects the metal

could have on the treatment and modelling within the planning system. The resultant

output factor was measured§ to be 0.06661 (± 0.00132) cGy/MU .

A detailed report accompanying the requested data was sent to the manufacturer.

§ On multiple occasions

57

C h a p t e r 3

DATA MODEL VALIDATION PHASE II

3 Dosimetric verification

In Phase II of this project beam data were sent to the manufacturer (Nucletron) for

source characterisation and the resultant beam model** was installed in the TPS at RBWH

for evaluation. Each model was to include the presence of the BUS. As detailed below it

was found necessary to seek refinements to this source model and as a consequence four

successive iterations of beam model were sent from the manufacturer for evaluation over

a period of 18 months, with each new beam model produced in response to shortcomings

identified in the previous one and reported to the manufacturer. Each of these beam

models allows selection of a pencil beam enhanced (PBE) and a collapsed cone enhanced

(CCE) dose calculation algorithm.

The dosimetric evaluation consists of evaluating the following parameters: absolute dose

output factor, the percentage depth dose at the treatment distance and heterogeneity

corrections under treatment conditions. These parameters were investigated and the

limitations of each beam model identified, using the methods outlined in this chapter. The

final model did not enable the calculation of dose using the collapsed cone convolution

algorithm; this deficiency precluded investigation of that component and it has been

acknowledged by the manufacturer as a known oversight.

3.1 Output Factor

Each treatment planning model includes a normalisation point, often referred to as the

calibration point at which an absolute output factor or calibration factor is assigned. As

stated in the previous chapter a calibration factor (0.06661 cGy/MU) was supplied to the

manufacturer as a part of the beam modelling process and is critical in assessing the

accuracy of the treatment planning system’s ability to calculate dose accurately. A plan was

generated in OMP that replicated the phantom and beam geometry and physical set-up

** Sent in the form of binary and database files

58

under which the calibration factor sent to the manufacturer was acquired (but ignoring the

presence of the BUS) and for each model the cGy/MU values calculated at the reference

point were extracted from the plan and recorded. The results are tabulated in the table 12.

Model 0 Model 1 Model 2 Model 3 Model 4

PBE 0.0638 0.0684 0.0684 0.0684 0.0684

CCE 0.0628 0.0685 0.0685 0.0686

Table 12: The dose at the reference point in cGy/MU calculated by the treatment planning system for each

beam model and calculation algorithm for a 40 cm x 25 cm field defined at the linac isocentre.

The original model (model 0) is seen to calculate a lower dose per monitor unit at the

calibration point than the other models. This occurs mainly because the model 0

calculation uses an output factor based on reference data acquired at 100 cm SSD rather

than the dose per MU factor supplied to the manufacturer for the re-characterisation and

does not account for the presence of the screen (as discussed in section 2.3). A consistent

discrepancy between the calculated dose per MU and the measured calibration value

(0.06661 cGy/MU) was noted for the re-characterised models and eventually after

consultation with the manufacturer it was discovered that during the data modelling

process the calibration factor supplied to the manufacturer was assigned to a 10 cm x 10

cm field (defined at 100 cm SSD) projected at a water phantom at 380 cm SSD, rather

than the 40 cm x 25 cm field that was used in the acquisition of the factor. This was

confirmed by repeating the above calculation using a 10 cm x 10 cm field in the treatment

planning system, the results of which are shown in table 13, where it is seen that the

discrepancy has been considerably reduced. For comparison, the discrepancies as a

percentage of the calibration value are tabulated in table 14.

Model 1 Model 2 Model 3 Model 4

PBE 0.0663 0.0663 0.0663 0.0661

CCE 0.0662 0.0662 0.0662

Table 13: The dose at the calibration point in cGy/MU calculated by the treatment planning system for each

beam model and calculation algorithm for a 10 cm x 10 cm field defined at the linac isocentre. The measured

calibration value was 0.06661 cGy/MU for a 40 cm x 25 cm field.

59

A consistent shift of approximately 2.8% should therefore be expected for all other TPS

calculations when comparing absolute doses as a result of this incorrect normalisation of

output, which is attributed to the change in collimator scatter (collimator scatter factor at

ESSD).

10 cm x 10 cm field 40 cm x 25 cm field

Model 1 Model 2 Model 3 Model 4 Model 1 Model 2 Model 3 Model 4

PBE -0.5% -0.5% -0.5% -0.8% 2.7% 2.7% 2.7% 2.7%

CCE -0.6% -0.6% -0.6% 2.8% 2.8% 3.0%

Table 14: Percentage difference between dose at the calculation point in cGy/MU calculated by the

treatment planning system for each beam model and calculation algorithm and the measured value of

0.06661 cGy/MU at the calibration point. A negative difference indicates a lower calculated dose than the

measured dose.

3.2 PDD

The percentage depth dose calculated by each treatment planning model was compared

with the PDD measured.

When each model was sent to RBWH for evaluation it was investigated for dosimetric

variations from measurements in the same fashion as described in chapter 2 of this thesis,

using the model to calculate doses for the same replicated phantom geometries as those

presented in chapter 2. Note that the replicated phantom geometries to which the models

were applied did not include the BUS, although it was present for the experimental

measurements. This was done intentionally because the screen cannot form part of the

CT images from which the patient geometry is derived and it is not practical to mimic the

screen as a bolus.

60

Percentage Depth Dose Comparision 7 cm below the Central Axis, Pencil Beam 40 cm x 25 cm 380 cm SSD

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se

Model 0

Model 1

Model 2

Model 3

Model 4

Measured

Figure 34: PDD as calculated at 7 cm below the CAX by the pencil beam algorithm for each treatment beam

model for a 40 cm x 25 cm field normalised to the maximum dose, in a large volume water phantom.

Note in the above graph in figure 34 the sudden dose drop after 50.25 cm depth for the

model 0 data. This is attributed to a limitation of the TPS data resulting from the

modelling process carried out by the manufacturer (standard practice excludes data

outside of 50.25 cm depth). The later models do not show this behaviour due to the

additional characterisation conducted by the manufacturer.

Percentage Depth Dose Comparision 7 cm below the Central Axis, Collapsed Cone 40 cm x 25 cm 380 cm SSD

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Dose

Model 0

Model 1

Model 2

Model 3

Model 4

Measured

Figure 35: PDD as calculated at 7 cm below the CAX by the collapsed cone algorithm for each treatment

beam model for a 40 cm x 25 cm field normalised to the maximum dose, in a large volume water phantom.

61

The dose distribution shown in the build-up region for models 2 and 3 as calculated by

the collapsed cone algorithm is not likely to be a true representation of the dose

deposition within this region. It is conjectured†† that the manufacturer has introduced low

energy photons or electrons into the beam model in an attempt to better calculate the

dose within the build-up region.

Percentage Depth Dose Comparision 43 cm Off Central Axis, Pencil Beam 40 cm x 25 cm 380 cm SSD

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se Measured

Model 1

Model 2

Model 3

Model 4

Figure 36: PDD as calculated at 7 cm below and 43 cm lateral to the CAX by the pencil beam algorithm for

each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose

†† As no detail of the modelling process conducted is provided by the manufacturer

62

Percentage Depth Dose Comparision 43 cm Off Central Axis, Collapsed Cone 40 cm x 25 cm 380 cm SSD

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se

Measured

Model 1

Model 2

Model 3

Figure 37: PDD as calculated at 7 cm below and 43 cm lateral to the CAX by the collapsed cone algorithm

for each treatment beam model for a 40 cm x 25 cm field normalised to the maximum dose

The results of gamma analysis and conclusions drawn for each model

3.2.1 Model 1


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 18.63 6.76 93.17

G_mean 2.58 4.45 6.38

Pass rate (%) 53.25 0.04 1.18

Table 15: The gamma analysis results comparing the measured PDD below the central axis and the PBE

Model 1 PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-


region (range 1 mm to 500 mm)

.

63


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 8.13 3.14 40.63

G_mean 2.05 2.96 3.27

Pass rate (%) 59.07 0.04 1.10

Table 16: The gamma analysis results comparing the measured PDD below the central axis and the CCE

Model 1 PDD with 10% or 2% and 1 mm distance to agreement, regions of interest include the In Build-Up

Region (IBUR), After Build-Up Region (ABUR) (range from dmax to 300 mm) and the user defined region

(range 1 mm to 500 mm)


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 19.92 6.18 65.03

G_mean 2.68 4.18 5.80

Pass rate (%) 53.25 0.04 1.18

Table 17: The gamma analysis results comparing the measured PDD 43 cm off axis and the PBE Model 1

PDD with 10% or 2% and 1 mm distance to agreement. Regions of interest include the In Build-Up Region


mm to 500 mm)


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 8.49 2.77 38.45

G_mean 2.09 1.92 1.70

Pass rate (%) 59.07 7.09 45.02

Table 18: The gamma analysis results comparing the measured PDD 43 cm off axis and the CCE Model 1



mm to 500 mm)

64

From the above comparison it was clear that the BUS had not been included in the

model. As a result the first model sent for evaluation (aka Model 1) was considered not

appropriate to be used clinically.

3.2.2 Model 2


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.20 4.07 6.74

G_mean 0.05 1.73 3.31

Pass rate (%) 100.00 40.87 25.35






Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 1.91 9.05 2.51

G_mean 1.91 0.40 0.24

Pass rate (%) 0.00 96.83 99.63






Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.32 4.17 6.52

G_mean 0.11 2.02 3.43

Pass rate (%) 100.00 28.26 18.21



65


mm to 500 mm)


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.55 10.04 10.04

G_mean 0.55 6.03 6.89

Pass rate (%) 100.00 0.23 0.04




mm to 500 mm)

From the above comparison although the new model (aka Model 2) was an improvement

on the previous one, with the dmax depth now coincident, there were obvious dosimetric

differences in the regions immediately following the build-up region which increased as a

function of depth from 20 cm. The CCE model was unsatisfactory because of the

anomalous dose distributions calculated by it in the build-up region. Thus upon evaluation

this data could not be used clinically and needed to be remodelled.

3.2.3 Model 3


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.42 4.07 6.75

G_mean 0.19 1.73 3.32

Pass rate (%) 100.00 41.07 24.94





66


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 1.96 9.29 9.29

G_mean 1.96 0.42 0.65

Pass rate (%) 0.00 94.34 79.64






Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.11 4.36 6.61

G_mean 0.04 2.20 3.61

Pass rate (%) 100.00 24.47 16.03




mm to 500 mm)


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 0.55 11.15 11.15

G_mean 0.55 6.95 8.09

Pass rate (%) 100.00 0.20 0.02




mm to 500 mm)

The visual comparison of PDDs for model 2 and model 3 revealed very similar curves,

leading to the conclusion that very minimal changes had been made by the manufacturer

67

to the energy fluence component in the characterisation process between these two

models, in which case both were insufficient to provide a viable clinical planning model.

3.2.4 Model 4


Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 2.42 2.37 12.10

G_mean 0.56 1.40 2.09

Pass rate (%) 77.48 28.61 17.54






Dist tol (mm) 1 1 1

Dose tol (%) 10 2 2

G_max 3.06 1.99 10.62

G_mean 0.63 1.35 1.80

Pass rate (%) 76.34 18.01 11.52




mm to 500 mm)

The build-up dose region was identified as having the poorest agreement; however given

that the majority of the dose errors occur in the calculation of dose in the lung much

deeper than 7 mm this was considered of limited clinical significance.

In consultation with the manufacturer it was determined that TPS energy characterisation

is conducted and validated by them for a 10 cm x 10 cm field. The data submitted for re-

characterisation was for a much larger (40 cm x 25 cm) field. This field size difference

results in an incorrect modelling of the beam dose variation with depth which was not

detected in the assessments made by the manufacturer and cannot be rectified by the user

68

once completed. If the user replicates a 10 cm x 10 cm field in OMP (conditions not

reflecting the experimental set-up), this results in a clinically acceptable agreement

between the measured (for the 40 cm x 25 cm field) and calculated (for the 10 cm x 10 cm

field) PDDs for models 2 and 3 as shown in figures 38 and 39. This observation was

relayed to the manufacturer and additional modelling was conducted to better replicate

the experimental set-up for model 4.

Percentage Depth Dose Comparision 7 cm below Central Axis, Pencil Beam, 10 cm x 10 cm 380 cm SSD, 40 cm x 25 cm

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se Measured

Model 1

Model 2

Model 3

Model 4

Figure 38: Treatment beam model PDD as calculated by the pencil beam algorithm for a 10 cm x 10 cm

field normalised at dmax rather than the 40 cm x 25 cm field, showing better agreement with the measured

PDD for the 40 cm x 25 cm field for models 2 and 3.

69

Percentage Depth Dose Comparision 7 cm below Central Axis, Collpased Cone, 10 cm x 10 cm 380 cm SSD, 40 cm x 25 cm

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se

Measured

Model 1

Model 2

Model 3

Figure 39: Treatment beam model PDD as calculated by the collapsed cone algorithm for a 10 cm x 10 cm

field normalised at dmax rather than the 40 cm x 25 cm field, showing better agreement with the measured

PDD for the 40 cm x 25 m field for models 2 and 3.

3.3 Profiles

The phantom geometry was replicated in the treatment planning system; the profiles

produced by each calculation algorithm were extracted and compared with the measured

data.

Profile Pencil Beam, Longitudinal profile, 400 cm Source to chamber distance, 40 x 25 cm field

25%

35%

45%

55%

65%

75%

85%

95%

105%

115%

-100 -80 -60 -40 -20 0 20 40 60 80 100

Position (cm)

Per

cen

tage

do

se Measured

Model 1

Model 2

Model 3

Model 4

Figure 40: Profiles as calculated at 7 cm below the CAX by the pencil beam algorithm for each treatment

beam model for a 40 cm x 25 cm field normalised to the central position

70

Profile Collapsed Cone, Longitudinal profile, 400 cm Source to chamber distance, 40 x 25 cm field

25%

35%

45%

55%

65%

75%

85%

95%

105%

115%

-100 -80 -60 -40 -20 0 20 40 60 80 100

Position (cm)

Per

cen

tage

do

se

Measured

Model 1

Model 2

Model 3

Figure 41: Profiles as calculated at 7 cm below the CAX by the collapsed cone algorithm for each treatment

beam model for a 40 cm x 25 cm field normalised to the central position.

It is observed in comparing the two figures 40 and 41 above, that in general the collapsed

cone algorithm yields profiles that better represent the measured data. The results of

gamma analysis of these beam profiles are tabulated in tables 29 - 35.

3.3.1 Model 1

Inner Penumbra Penumbra

Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 1.49 1.32 1.36

G_mean 0.27 0.35 0.33

Pass rate (%) 96.18 94.19 98.81


Table 29: The results of gamma analysis performed to compare the Pencil Beam calculation of the profile

and the measured data for Model 1.

71


Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 0.67 1.82 8.43

G_mean 0.16 0.31 0.55

Pass rate (%) 100.00 98.84 92.68




3.3.2 Model 2


Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 1.55 1.13 0.84

G_mean 0.30 0.35 0.31

Pass rate (%) 95.42 93.02 100.00





Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 0.75 1.64 6.86

G_mean 0.16 0.31 0.45

Pass rate (%) 100.00 98.84 93.90




72

3.3.3 Model 3


Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 1.54 1.16 0.84

G_mean 0.30 0.35 0.31

Pass rate (%) 95.49 93.02 100.00





Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 0.73 1.67 6.98

G_mean 0.16 0.31 0.46

Pass rate (%) 100.00 98.84 93.90




3.3.4 Model 4


Left Right

Dist tol (mm) 2 3 3

Dose tol (%) 2 3 3

G_max 0.96 0.75 0.50

G_mean 0.17 0.32 0.25

Pass rate (%) 100.00 100.00 100.00




73

In general, the inner region of the profile produced by each beam model yielded

acceptable agreement for the collapsed cone calculations. The regions colloquially known

as the “horns” in each profile demonstrate poorest agreement with the measured data.

This region will vary greatly with small changes in the beam spectra. Consequently the

final model yields the best agreement with the measured profile data, because the energy

of model 4 best represents that of the measured beam, as evidenced by the superior

agreement of the PDD produced (discussed in section 3.2.4). As discussed in section 2.6

the significance of the profile calculation error is limited, since the beam will in most

clinical cases overshoot the patient at the location of the horns.

3.4 Simple phantom comparison

As described in section 2.8 of this thesis, a Farmer chamber was placed at the mid-point

of a solid water phantom, the thickness of which was then varied and doses recorded for a

delivered 400 MU for each thickness. The simple solid water phantom geometries were

replicated in the TPS and dose calculations performed by both pencil beam (PBE) and

collapsed cone (CCE) algorithms for all beam models. The doses calculated at each

chamber depth by the TPS for 400 MU were compared with the measured doses and the

percentage differences are tabulated.

PBE CCE

Depth Model 0 Model 1 Model 2 Model 3 Model 4 Model 0 Model 1 Model 2 Model 3

5 7.3% 14.7% 14.7% 14.7% 16.4% 4.8% 14.1% 14.1% 14.2%

14 8.6% 16.2% 16.2% 16.2% 13.9% 5.2% 14.7% 14.7% 14.7%

16 8.7% 16.4% 16.4% 16.4% 13.2% 5.8% 15.4% 15.4% 15.3%

21 9.1% 16.7% 16.7% 16.7% 11.5% 5.3% 15.0% 15.0% 14.8%

26 8.5% 16.1% 16.1% 16.1% 9.3% 6.6% 16.3% 16.3% 15.9%

Table 36: The percentage differences between the doses measured in the simple solid water phantom and

those calculated using the PBE and CCE algorithms for all beam models investigated for the same phantom

geometry and number of MU in OMP. The positive differences indicate higher calculated doses than the

measured doses.

For the PBE models, the model 0 discrepancy does vary slightly with thickness with an

average deviation from measured doses of 8.4 %. Models 1 – 3 show identical

discrepancies the average of which is 16 %, with the pattern of variation with depth the

74

same as for model 0. Model 4 shows the largest variation in calculation discrepancy, with

the measured values, ranging from 9.3 % to 16.4 %.

In comparing the ratio of the doses calculated by each re-characterised model to the

original model, very limited change in this ratio is observed for all separations considered.

It is inferred from this internal consistency that there is no significant change in the

convolution kernel between re-characterised models 1 - 3. Models 1 and 2 procedure the

same deviation from the measured data and for accurate dose calculations in this case a

scaling (or output) factor can be applied to yield reasonably accurate dose calculations

since the variations are also fairly independent of depth.

3.5 Measurement consistency investigation

The deviation of the TPS calculated dose from measured data in the case of the absolute

output factor comparison in section 3.1 was observed to be small when compared with

the deviation observed in section 3.4 for the small simple solid water phantom. A

comparison of the two data sets was made to try to determine the cause of this

inconsistency. In order to see whether the two sets of measured data appeared to be

consistent with each other an attempt was made to derive the expected measured value in

the small simple phantom from the measured dose value in the large absolute dosimetry

phantom by applying reasonable corrections based on the inverse square law and the

effects of phantom size on phantom scatter.

From the measured data from section 2.8 an interpolated measured dose at 10 cm depth

in the simple phantom is found to be 21.9 cGy for 400 MU.

An expected measured dose in the simple phantom can be calculated from the dose

measured in the absolute phantom case, assuming that the change in the scattering

volume has limited effect and the inverse square law applies.

Dose from the water phantom for 400 MU was 26.64 cGy. This measurement was

acquired at a SCD of 390 cm so to convert this dose to an equivalent dose for the

measurement at 400 cm SCD the inverse square law is applied:

75

cGy 32.25400390

*64.262

=

The expected change in measured dose due to the phantom scatter can be accounted for

by comparing the PDDs at 10 cm depth for the areas presented to the beam by the two

phantoms, using standard tables of PDDs (acquired at 100 cm SSD) as a function of field

size.

For the absolute dose phantom the area presented to the beam was 1800 cm2. The PDD

at 10 cm for the 10 MV beam is 72.6 % for this area.

For the simple phantom the area presented to the beam was 400 cm2. The tabulated PDD

at 10 cm for the 10 MV beam is 75.2 %.

This leads to an expected increase in dose of ~ 3.5% due to the increased phantom scatter

in the larger phantom. The change in the backscatter would also contribute, but this is

much less than the calculated 3.5% change which is predominantly due to forward scatter.

Applying this correction the expected dose is:

cGy 44.24965.0*32.25 =

Therefore the expected dose can be directly compared with the interpolated measured

dose

116.19.21

44.24 =

The difference between the two doses, one directly measured and the other inferred from

absolute dose (output factor) measurements is 11.6%. Taking into account the known

error in the normalisation of the absolute output factor for models 1 to 4 (which is of the

order of 2.9%), the observed discrepancy between TPS calculated dose and measured

dose of ~ 14% for the simple phantom is consistent with this difference. It was concluded

that there may an error in the output factor supplied to the manufacturer and

76

incorporated in models 1 – 4, but to confirm this, the measurements from which the

output factor was derived will need to be repeated.

3.6 Thorax Phantom Comparison

As described in section 2.10, dose measurements in the lung cavity closest to the beam,

the central soft tissue cavity and the spinal bone cavity were acquired for 500 MU. The

thorax phantom CT data set used in section 2.11 was again employed to compare dose

calculations performed by both pencil beam (PBE) and collapsed cone (CCE) algorithms

for all beam models to the measured data. The dose for comparison with the TPS

calculated dose is dose to water within these bone, tissue and lung regions. The Pinpoint

chamber used has a very low collecting volume and hence the effect of cable signal

becomes significant for the length of cable irradiated, increasing the uncertainty in the

readings (and thus the doses) in the thorax phantom case. The extra cable signal was

measured for the Pinpoint chamber; yielding 7.1% extra signal induced for the

approximate length of cable present in the experimental set-up.

PBE

Region Model 0 Model 1 Model 2 Model 3 Model 4

Lung -2.9% 3.8% 3.8% 3.8% 4.1%

Bone -9.1% -2.9% -2.9% -2.9% -7.0%

Tissue -10.5% -4.3% -4.3% -4.3% -8.6%

CCE

Region Model 0 Model 1 Model 2 Model 3

Lung -5.7% 2.8% 2.8% 2.9%

Bone -3.2% 5.6% 5.6% 5.5%

Tissue -4.3% 4.5% 4.5% 4.5%

Table 37: The percentage difference between doses measured by the ionisation chamber in the tissue, lung

and bone and those calculated by the TPS for 500 MU, where the chamber cavity has been replicated in the

TPS as being filled with water. The ionisation chamber readings have been corrected for cable signal. A

negative difference indicates lower calculated dose than the measured dose.

In spite of the variation in the dose with depth (when normalised to a depth of dmax)

between models 1, 2 and 3, observed in section 3.2, there is very limited variation in the

calculated doses between the first three models within the thorax phantom, for the PBE

and CCE algorithms individually.

77

This internal consistency implies that there is little difference in the convolution kernels

employed in the three models, and may be enhanced by the fact that the models are all

designed to deliver a particular dose per MU at 10 cm depth in water, which is

approximately the water equivalent depth of the thorax phantom measurements. The

similarity between models for PDDs normalised at 10 cm depth in water is shown in the

figures 42 and 43.

Percentage Depth Dose Comparision 7 cm below the Central Axis, Pencil Beam 40 cm x 25 cm 380 cm SSD

0%

20%

40%

60%

80%

100%

120%

140%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

cen

tage

Do

se

Model 0

Model 1

Model 2

Model 3

Model 4

Measured

Figure 42: PDD as calculated by the pencil beam algorithm for each treatment beam model for a 40 cm x 25

cm field normalised to the depth of 10 cm.

78

Percentage Depth Dose Comparision 7 cm below Central Axis, Collapsed Cone 40 cm x 25 cm 380 cm SSD

0%

10%

20%

30%

40%

50%

60%

70%

80%

90%

100%

110%

120%

130%

140%

150%

0 5 10 15 20 25 30 35 40 45 50 55

Depth (cm)

Per

centa

ge D

ose Model 0

Model 1

Model 2

Model 3

Measured

Figure 43: PDD as calculated by the collapsed cone algorithm for each treatment beam model for a 40 cm x

25 cm field normalised to the depth of 10 cm.

3.7 Anthropomorphic Phantom

This investigation compares the doses calculated by both algorithms of each beam model

at selected points in the anthropomorphic phantom with the doses measured previously

with TLD at those same points as described in sections 2.12 and 2.13. This was done to

evaluate the performance of the algorithms in calculating dose in the presence of

heterogeneities in a phantom system more closely representing a patient than the simple

water phantom or thorax phantom. The results of this comparison are included in tables

41 and 42 in the appendix.

79

Anthropomorphic Phantom PBE dose calcuation variation

-20.0%

-15.0%

-10.0%

-5.0%

0.0%

5.0%

10.0%

15.0%

20.0%

25.0%

30.0%


Version of beam model

Per

cen

tage

dif

fere

nce

fro

m m

easu

rem

ents

Average

max

min

Figure 44: The graph displaying the average, minimum and maximum percentage deviations of the TPS

calculated doses from the TLD doses for Pencil Beam calculations.

The graph shown in figure 44 demonstrates the degree of overall agreement between the

PBE model calculations and the TLD measurements. The spread between the maximum

and minimum deviations is least for models 2 and 3, which show identical values for all

three parameters plot. This is consistent with the similarities in the PDDs noted in section

3.2 for these models.

80

Anthropomorphic phantom CCE dose calcualtion variation

-15.0%

-10.0%

-5.0%

0.0%

5.0%

10.0%

15.0%

20.0%

25.0%

Model 0 Model 1 Model 2 Model 3

Version of beam model

Per

centa

ge d

iffe

rence

fro

m m

easu

rem

ents

average

min

max

Figure 45: The graph displaying the average percentage deviations of the TPS calculated doses from the

TLD doses for Collapsed Cone calculations.

The graph shown in figure 45 demonstrates the degree of overall agreement between the

CCE model calculations and the TLD measurements. The average TPS calculated dose

deviation from the measured TLD dose is higher by ~ 10% for each of the re-

characterised models, differing only slightly from the corresponding average for the PBE

models. Although this can partly be attributed to the incorrect handling of the calibration

factor conditions within the TPS as discussed in section 3.1 of this thesis, a large residual

deviation remains, which is consistent with the postulated error in the output factor

mentioned previously.

3.7.1 TLD variation investigation

An anomalous result was observed in the TLD measurements in the anthropomorphic

phantom, viz. that the TPS calculated doses were higher than the corresponding TLD-

measured doses by up to 20%. The discrepancies are much greater than the uncertainty

associated with the TLD measurements (discussed in chapter 1). The magnitude of the

type A uncertainty associated with the TLD measurements can be ascertained from the

81

standard deviation of the mean of the 4 TLD measurements at each location in the

phantom.

To investigate this anomaly it was decided to compare doses measured using TLDs with

those measured using an ion chamber at the same point in a composite

anthropomorphic/ solid water phantom at the ESSD (see table 38). This procedure

provides an alternative calibration of the TLDs at the ESSD to their standard calibration

at 100 cm SSD.

Several slices of the anthropomorphic phantom were replaced by the solid water slabs (as

shown in figure 46).

Figure 46: The anthropomorphic/ solid water phantom arrangement used in the TLD accuracy and linearity

investigation.

3.7.1.1 TLD calibration at ESSD

An ion chamber‡‡ was placed in the mid-sagittal plane at the central axis within the

phantom shown in figure 46, at a SCD of 400 cm. Charge was collected and converted to

dose for four different MU values. The ion chamber was replaced with TLDs which

irradiated using 100, 250, 500 and 1000 MU. The doses in cGy obtained for both types of

‡‡ Nuclear Enterprises(Nuclear Enterprises, Reading, England) 2571 Farmer type chamber

82

dosimeter are found in table 38. The TLD doses are derived using their standard

calibration at 100 cm SSD.

MU Chamber

doses Stdev

Chamber

Chamber

(uA) TLD doses

TLD stdev

TLD

(uA) Chamber /

TLD

100 5.5 0.07 0.04 5.1 0.05 0.02 7%

250 13.5 0.02 0.01 12.9 0.08 0.03 5%

500 27.1 0.00 0.00 25.9 0.32 0.15 5%

1000 54.1 0.13 0.07 51.4 0.42 0.19 5%

Table 38: The results of the comparison of TLD dose measurements in the anthropomorphic/ solid water

phantom with doses measured by the ionisation chamber.

These results appear to demonstrate an under-response of the TLDs when used for dose

measurements in a large volume at ESSD, which leads to an average 5 % deviation from

ionisation chamber measured dose. The reasons for this 5% discrepancy are unknown,

and require further investigation. However the TLDs were assigned new calibration

factors based on the ESSD re-calibration, and the anthropomorphic phantom results were

revised and are tabulated in tables 43 and 44 in the appendix.

Although this correction improves the overall agreement (the differences between the

minimum and maximum in addition to the general average agreement) of the TPS dose

calculations with the TLD results, differences still remain for all versions of the beam

model. The largest variation between the TPS-calculated doses and measurements occurs

within the build-up region.

The percentage deviations of the TPS calculated doses for model 0 from the revised TLD

doses taken from tables 43 and 44 are shown in relation to the locations in the phantom

where they appear in figures 47.1 – 47.10.

84

Figures 47.1 – 47.10: CT slices of the anthropomorphic phantom, showing the individual TLD positions

within the phantom. The percentage differences from the TLD of doses calculated by collapsed cone and

pencil beam algorithms are shown in blue and red respectively for model 0. The TLDs were calibrated in a

large phantom at ESSD.

85

Corrected TLD comparision PBE

-20.0%

-15.0%

-10.0%

-5.0%

0.0%

5.0%

10.0%

15.0%

20.0%


average

min

max


calculated doses from the recalibrated TLD doses for PBE calculations.

TLD corrected CCE comparision

-20.0%

-15.0%

-10.0%

-5.0%

0.0%

5.0%

10.0%

15.0%

20.0%

model 0 model 1 model 2 model 3

average

min

max


calculated doses from the recalibrated TLD doses for CCE calculations.

86

C h a p t e r 4

DISCUSSION AND CONCLUSION

…….

4 Project Overview

The eventual phasing out of the PLATO 2D treatment planning system currently used to

plan TBI treatments at RBWH and the acquisition of a more modern and sophisticated

TPS, Oncentra Masterplan (OMP), generated the need to commission OMP for TBI

treatment planning. Two treatment planning calculation algorithms (a pencil beam and a

collapsed cone) and a total of five separate beam models were evaluated for their ability to

calculate TBI doses with sufficient accuracy for clinical implementation.

The sources of uncertainty in any radiotherapy treatment delivery are discussed by Van

Dyk[14]. This thesis aimed to quantify the “beam measurement uncertainty” by

measurement of the treatment beam properties such as output, profiles and PDDs, and to

quantify the accuracy associated with dose calculations by comparing the dose calculations

from OMP with measurements under TBI conditions. A series of TLD measurements

acquired within an anthropomorphic phantom was used to quantify the “patient related”

uncertainty in reproducibility of set-up and to assess the ability of OMP to handle

heterogeneities. The “mechanical and dosimetric treatment machine-, imaging- related

uncertainties, dose evaluation uncertainties, biological modelling uncertainties” have been

previously quantified for 100 cm SAD treatments at RBWH and are not within the scope

of this thesis. The uncertainty associated with the target and normal tissue delineation is

not considered for this thesis as the entire body is regarded as the target and the normal

tissue delineation accuracy is not considered to be of critical importance in this treatment

process.

The results of the various comparisons of calculated and measured doses are summarised

and discussed in the following section.

87

4.1 Discussion

4.1.1 PDDs

PDDs measured close to and well away from the beam axis were compared to PDDs

calculated by the treatment planning system. Gamma analysis was carried out and revealed

clinically unacceptable deviations for the OMP beam model already implemented in the

TPS for conventional treatments, using both Pencil Beam and Collapsed Cone algorithms.

Following the re-characterisation of the photon fluence, the PDDs produced by each new

beam model were also compared with those measured, and again the results were not

clinically acceptable. This was partly because the TPS was not capable of replicating the

dosimetric effects of the BUS used in TBI.

An arbitrary restriction of the calculation of doses using PBE for model 0 is observed in

the calculated PDD for depths greater than 50.25 cm. Comparison of the plots of the

subsequent models with the measured data revealed that the CCE algorithm

demonstrated consistently poorer agreement within the build-up region. Conversely the

PBE algorithm demonstrated a diminished agreement for depths greater than 20 cm. It

was subsequently discovered that for all re-characterised models, the PDD data that was

supplied to the manufacturer was parameterised by them on the assumption that it was

for a 10 cm x 10 cm field rather than the 40 cm x 25 cm field used to acquire the data.

This would adversely affect the agreement with the measured data for models 1 to 4.

4.1.2 Beam profiles measured in air

A profile acquired in air was compared with profiles for each beam model calculated by

CCE and PBE. There were no significant variations between the models or algorithms

with the exception of model 4 which demonstrated slightly better agreement with the

measured data. Regions of poorest agreement fall outside the bounds of patient geometry

and this discrepancy is not clinically significant.

4.1.3 Output factor

The reference output factor at 10 cm depth was required to be supplied as a factor for the

beam re-characterisation in the generation of beam models. This data was acquired for a

field size set to 40 cm x 25 cm. However this data was included in these beam models as if

88

it were for a 10 cm x 10 cm field size. This results in an overestimation of about 3% in the

dose calculation for fields replicated in the TPS with models 1 to 4. A dose overestimation

of about 3% would be expected for model 0 beyond the build-up region due to the

presence of the BUS during measurements and its absence in the TPS calculations.

4.1.4 Thorax phantom

Measurements with a thorax phantom were used to quantify doses within regions of

heterogeneity. Calculations based on models 1 to 4 exhibited similar discrepancies

between the calculated and measured doses. This similarity is not unexpected since the

points of comparison lie in the post build-up region where the differences between beam

models are smaller than in the build-up region for each beam model.

For models 1 – 4 for PBE the doses calculated were lower than those measured in the

bone and soft tissue analogues, but higher within the lung, whereas for models 1 – 3 the

CCE algorithm demonstrated a higher dose calculated than measured in all three tissue

types.

4.1.5 Simple phantom

The phantom dose measurements were made in the mid-sagittal region over a range of

phantom thicknesses. This was done to evaluate the accuracy of the dose calculations over

a range of patient separations. Doses calculated by models 1 - 3 for both PBE and CCE

exhibited similar discrepancies with the measured data for each separation, which were

above the level of clinical acceptability. Model 4 varied slightly from the previous models

considered, however the discrepancies still fell outside the level of clinical acceptability of

5%.

4.1.6 Anthropomorphic phantom

Two sets of data were acquired using TLDs placed inside an anthropomorphic phantom

to investigate the three-dimensional dose distribution in a phantom plane. The calculated

doses in the bulk of the phantom volume are consistently higher than measured but

within clinically acceptable ranges. In the vicinity of the build-up region the agreement is

much poorer as expected from results of section 3.2. These results show that the beam

models could be accommodated into a clinical protocol in which the limitations of the

89

dose calculations in the superficial regions are recognised. However the concern of exit

dose and build-up is reduced in using an opposed beam configuration as is the case for

the proposed treatment technique.

Disagreements between the measured and calculated doses may be attributable to

limitations in the models’ ability to accurately account for scatter (air, BUS and backscatter

from the concrete wall). The significance of this is discussed in work by Poole[74] using

Monte Carlo calculations. That work demonstrates the increased importance of

appropriately replicating the details of the scattering environment for dose calculations at

the large distances appropriate for TBI.

Note that the simulated phantom geometries to which the models were applied did not

include the BUS, although it was present for the actual dose measurements. This was

done intentionally because the screen cannot form part of the CT images from which the

patient geometry is derived and it was desired to avoid the added complication of adding

it to the plan subsequently as a bolus with a large air gap to the patient.

4.1.7 General Discussion

During the course of this investigation, OMP released some minor updates (V4.1 SP2)

and some major updates (V3.2) of the planning software. For each version of the

treatment planning system, the methods of comparison presented in this thesis were made

at the time (for all available models). In total five separate versions of OMP have been

investigated (V3.1 SP2, V3.2, V3.3, V4.0 and V4.1 SP2). Only results of the final

investigation (V4.1 SP2) are presented in this thesis. A note of interest: OMP originally

made four algorithms available for calculation Pencil Beam Classic, Pencil Beam

Enhanced, Collapsed Cone Classic and Collapse Cone Enhanced; however after the

release V4.0 of OMP the “Classic” algorithms are no longer available. The results of these

algorithms are no longer relevant to the clinical implementation of TBI at RBWH and are

not included in this thesis. No detectable variations in the calculated TBI doses for PBE

and CCE were observed between updates of the TPS software.

A disadvantage in the implementation of a collapsed cone calculation algorithm compared

to a pencil beam algorithm is the increase in calculation time. For all versions of the

90

treatment planning system investigated a calculation time comparison was carried out. In

every case a longer calculation time was observed for the collapsed cone (calculation times

generally of the order of 90 s) than for the pencil beam (calculation times generally of the

order of 40 s). This difference was not sufficient to influence the choice of algorithm in

the clinical implementation of TBI given that the plans will be limited to opposed parallel

pairs.

4.2 Conclusion and Future Work

Four refinements of the original beam model were implemented in the TPS. Calculations

based on the fourth and most recent model are most consistent with the measured data

beyond the region of build-up. However the build-up region agreement was much poorer

attributed to the incomplete characterisation of the effect of the BUS. Thus of the five

models investigated, model 4 appears to be the best candidate for clinical implementation.

The output factor would need to be adjusted to bring calculated absolute doses beyond

the build-up region into line with experiment, and the reduced accuracy of the calculated

doses in the build-up region would need to be accepted. However it is possible that the

theoretically superior CCE version of model 4 will give even better results, and a final

decision on clinical implementation has been withheld pending its availability and

evaluation.

TBI treatment planning based on OMP at the RBWH has not been proven to be clinically

acceptable. It is believed OMP in fully utilising the patient CT data should predict dose

distributions when bone and lung are included in the calculations more accurately than the

existing protocol. More work must be undertaken in patient specific dosimetry of these

treatments through the use of in-vivo dosimetry with thermo- and optically-stimulated

luminescent dosimetry as well as ionisation chambers for the clinical implementation of

TBI treatments. Although this investigation has aimed to wholly complete the

commissioning process, none of the models examined yielded dose calculations that were

within the clinical acceptance range of 5% agreement with measured dose distributions for

all conditions. RBWH is still in consultation with the manufacturer to rectify the gross

errors observed in the calculation.

91

Once a clinically acceptable model is commissioned, future project work will include the

feasibility of utilising virtual bolus or real bolus at the simulation appointment for the

compensation for lungs and significant contour variations (as in cases of treatment of

paediatric patients). An additional investigation required is the accuracy of modelling the

effects of lead shielding in OMP for use as compensators to reduce dose to organs such as

the brain. The use of bolus material has already been implemented clinically for

conventional radiotherapy in OMP and its use will be investigated for TBI. The

transmission through the bolus materials to be used will be quantified and compared to

that calculated in the TPS.

The inconsistency in the deviation of the calculated doses from the measured doses was

briefly investigated, pointing to an inconsistency between measured absolute doses using

different measuring systems. The source of this error is unknown: the depth at which the

output factor for models 1 – 4 has been normalised needs to be clarified and RBWH is in

consultation with the manufacturer to resolve this issue. In addition an apparent change in

response of the TLDs within the large phantom at ESSD was observed; the cause of this

variation is unknown and will be investigated through repetition of experimental data

acquisition.

There is scope to verify this data with Monte Carlo (BEAMnrc) calculations, the feasibility

of which was investigated by Poole[74], to enable a more robust investigation into TBI

dosimetry. The RBWH has recently developed a high performance computer cluster to

enable the use of Monte Carlo dosimetry verification for clinical treatments and its use in

TBI dosimetry is deemed a priority and will be implemented.

92

A p p e n d i x

APPENDIX

5 Appendices

5.1 Phase 1 Pencil Beam TLD Results

PBE

Dose point Dose % diff TLD

1 61.9 10.4%

2 66.4 11.3%

3 73.1 9.3%

4 60.2 12.9%

5 71.2 8.5%

6 73.7 7.9%

7 62.4 11.4%

8 37.8 10.2%

9 36.8 13.3%

10 48.5 9.6%

11 57.2 11.5%

12 63.9 13.3%

13 73.2 -1.9%

14 74.9 2.7%

15 42.6 6.4%

16 51.3 9.0%

17 60.4 4.6%

18 68.5 5.3%

19 68.3 -9.1%

20 42.9 7.1%

21 39.5 12.4%

22 71.0 15.4%

23 47.8 8.0%

24 57.3 6.9%

25 52.2 6.2%

26 43.4 7.3%

27 63.0 5.3%

28 47.6 5.6%

29 47.9 7.8%

30 49.4 6.0%

31 68.2 5.1%

32 70.3 6.5%

33 38.3 7.9%

34 46.6 7.1%

35 55.9 4.8%

93

36 66.2 3.3%

37 63.5 -13.3%

38 62.5 5.2%

39 58.1 1.5%

40 66.1 9.5%

41 50.9 14.4%

average 7.0%

min -13.3%

max 15.4%

Table 39: The doses in cGy measured with TLDs within the anthropomorphic phantom and comparison

with the existing beam model (model 0) doses calculated by the Pencil Beam algorithm in OMP. The “% diff

TLD” is the percentage by which the calculated dose exceeds the measured dose.

5.2 Phase 1 Collapsed Cone TLD Results

CCE

Dose point Dose % diff TLD

1 59.9 6.7%

2 64.7 8.3%

3 72.4 8.1%

4 57.4 7.6%

5 69.8 6.5%

6 73.4 7.5%

7 60.7 8.4%

8 36.0 4.9%

9 35.2 8.5%

10 46.5 5.0%

11 55.2 7.7%

12 60.8 7.8%

13 71.1 -4.8%

14 72.7 -0.3%

15 41.4 3.4%

16 49.6 5.3%

17 59.7 3.5%

18 65.9 1.4%

19 67.6 -10.0%

20 38.7 -3.4%

21 35.9 2.4%

22 67.3 9.4%

23 45.9 3.7%

24 55.4 3.2%

25 50.3 2.3%

26 41.3 2.2%

27 63.0 5.3%

28 46.0 2.0%

94

29 46.5 4.6%

30 47.4 1.7%

31 66.0 1.6%

32 67.9 2.9%

33 36.2 1.9%

34 44.6 2.6%

35 53.9 1.0%

36 64.8 1.2%

37 64.7 -11.7%

38 62.3 4.9%

39 57.1 -0.4%

40 63.2 4.7%

41 48.4 8.7%

average 3.3%

min -11.7%

max 9.4%

Table 40: The doses in cGy measured with TLDs within the anthropomorphic phantom and comparison

with the existing beam model (model 0) doses calculated by the Collapsed Cone algorithm in OMP. The “%

diff TLD” is the percentage by which the calculated dose exceeds the measured dose.

95

5.3 Anthropomorphic phantom TLD comparison

dose stdev st uncert1 61.9 10.4% 66.3 18.2% 66.3 18.2% 66.3 18.2% 66.4 18.3% 56.09 2.23 1.112 66.4 11.3% 71.1 19.2% 71.1 19.2% 71.1 19.2% 72.1 20.8% 59.68 0.92 0.463 73.1 9.3% 78.1 16.7% 78.1 16.7% 78.1 16.7% 80.4 20.1% 66.92 3.71 1.854 60.2 12.9% 64.4 21.0% 64.4 21.0% 64.4 21.0% 64.5 21.0% 53.28 0.10 0.055 71.2 8.5% 76.1 16.1% 76.1 16.1% 76.1 16.1% 77.9 18.8% 65.58 2.42 1.216 73.7 7.9% 78.8 15.4% 79.4 16.3% 79.3 16.2% 81.0 18.6% 68.29 1.32 0.667 62.4 11.4% 66.8 19.3% 66.8 19.3% 66.8 19.3% 66.8 19.3% 56.01 0.95 0.488 37.8 10.2% 40.4 17.7% 40.4 17.7% 40.4 17.7% 37.8 10.2% 34.29 0.16 0.089 36.8 13.3% 39.3 21.0% 39.3 21.0% 39.3 20.9% 36.7 13.1% 32.47 0.50 0.2510 48.5 9.6% 51.9 17.3% 51.9 17.3% 51.9 17.3% 49.5 11.9% 44.24 0.54 0.2711 57.2 11.5% 61.1 19.2% 61.1 19.2% 61.1 19.2% 59.5 16.1% 51.28 0.72 0.3612 63.9 13.3% 68.3 21.2% 68.3 21.2% 68.3 21.2% 68.1 20.8% 56.37 1.85 0.9313 73.2 -1.9% 78.5 5.0% 88.6 18.6% 88.2 18.0% 88.7 18.8% 74.69 0.63 0.3214 74.9 2.7% 80.3 10.1% 88.5 21.4% 88.1 20.9% 89.0 22.0% 72.92 0.77 0.3915 42.6 6.4% 45.5 13.6% 45.5 13.6% 45.5 13.6% 42.8 6.9% 40.02 0.41 0.2016 51.3 9.0% 54.9 16.6% 54.9 16.6% 54.9 16.6% 52.9 12.4% 47.07 1.49 0.7417 60.4 4.6% 64.6 11.9% 64.6 11.9% 64.6 11.9% 63.9 10.6% 57.72 0.47 0.2418 68.5 5.3% 73.3 12.7% 73.3 12.7% 73.3 12.7% 74.5 14.6% 65.03 2.01 1.0019 68.3 -9.1% 73.1 -2.8% 86.2 14.7% 85.9 14.2% 86.1 14.5% 75.18 1.38 0.6920 42.9 7.1% 45.9 14.6% 45.9 14.6% 45.9 14.6% 43.6 8.8% 40.08 0.97 0.4821 39.5 12.4% 42.2 20.1% 42.2 20.1% 42.2 20.1% 39.6 12.8% 35.12 0.09 0.0422 71.0 15.4% 76.0 23.5% 76.0 23.5% 76.0 23.5% 76.9 25.0% 61.56 1.65 0.8223 47.8 8.0% 51.2 15.7% 51.2 15.7% 51.2 15.7% 49.8 12.6% 44.25 1.17 0.5824 57.3 6.9% 61.4 14.5% 61.4 14.5% 61.4 14.5% 61.0 13.7% 53.64 1.37 0.6925 52.2 6.2% 56.0 13.8% 56.0 13.8% 56.0 13.8% 55.0 11.8% 49.20 0.37 0.1926 43.4 7.3% 46.5 14.9% 46.5 14.9% 46.5 14.9% 44.7 10.5% 40.44 0.77 0.3827 63.0 5.3% 67.0 12.0% 67.0 12.0% 67.0 12.0% 67.0 12.0% 59.82 0.54 0.2728 47.6 5.6% 50.9 13.0% 50.9 13.0% 50.9 13.0% 48.5 7.6% 45.06 0.74 0.3729 47.9 7.8% 51.3 15.4% 51.3 15.4% 51.3 15.4% 48.9 10.1% 44.44 1.28 0.6430 49.4 6.0% 52.9 13.4% 52.9 13.4% 52.9 13.4% 50.8 8.9% 46.64 2.35 1.1731 68.2 5.1% 73.0 12.5% 73.0 12.5% 73.0 12.5% 74.1 14.1% 64.93 2.55 1.2832 70.3 6.5% 75.2 13.9% 75.2 13.9% 75.2 13.9% 76.5 15.8% 66.04 4.03 2.0233 38.3 7.9% 41.0 15.5% 41.0 15.5% 41.0 15.4% 38.9 9.6% 35.49 0.28 0.1434 46.6 7.1% 49.9 14.7% 49.9 14.7% 49.9 14.7% 48.2 10.8% 43.49 1.08 0.5435 55.9 4.8% 59.8 12.2% 59.8 12.2% 59.8 12.2% 58.9 10.5% 53.32 1.35 0.6736 66.2 3.3% 70.9 10.6% 70.9 10.6% 70.9 10.6% 71.3 11.2% 64.10 2.18 1.0937 63.5 -13.3% 67.9 -7.3% 83.4 13.8% 83.1 13.4% 82.3 12.3% 73.29 0.27 0.1338 62.5 5.2% 66.9 12.6% 66.9 12.6% 66.9 12.6% 66.8 12.5% 59.40 1.06 0.5339 58.1 1.5% 62.2 8.5% 62.2 8.5% 62.2 8.5% 60.9 6.3% 57.28 2.19 1.0940 66.1 9.5% 70.8 17.2% 70.8 17.2% 70.8 17.2% 71.2 18.0% 60.35 0.36 0.1841 50.9 14.4% 54.4 22.4% 54.4 22.4% 54.4 22.4% 52.4 17.8% 44.48 0.11 0.06

average 7.0% average 14.5% average 16.0% average 16.0% average 14.2%min -13.3% min -7.3% min 8.5% min 8.5% min 6.3%max 15.4% max 23.5% max 23.5% max 23.5% max 25.0%

TLDsPBEModel 0 Model 1 Model 2 Model 3 Model 4

Table 41: The doses calculated using the PBE algorithm for each beam model, at each TLD position, and a

comparison with the doses measured by the TLDs. For each model, the first column shows the calculated

dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the

TLD-measured dose at that point for the same number of MU.

96

dose stdev st uncert1 59.9 6.7% 65.3 16.4% 65.3 16.4% 65.3 16.4% 56.09 2.23 1.112 64.7 8.3% 70.5 18.2% 70.5 18.2% 70.6 18.3% 59.68 0.92 0.463 72.4 8.1% 78.7 17.6% 78.7 17.6% 78.8 17.7% 66.92 3.71 1.854 57.4 7.6% 62.6 17.4% 62.6 17.4% 62.6 17.4% 53.28 0.10 0.055 69.8 6.5% 76.1 16.0% 76.1 16.0% 76.1 16.1% 65.58 2.42 1.216 73.4 7.5% 80.1 17.2% 80.7 18.1% 80.7 18.1% 68.29 1.32 0.667 60.7 8.4% 66.3 18.3% 66.3 18.3% 66.3 18.3% 56.01 0.95 0.488 36.0 4.9% 39.3 14.6% 39.3 14.6% 39.2 14.2% 34.29 0.16 0.089 35.2 8.5% 38.5 18.5% 38.5 18.5% 38.3 18.1% 32.47 0.50 0.25

10 46.5 5.0% 50.7 14.6% 50.7 14.6% 50.7 14.5% 44.24 0.54 0.2711 55.2 7.7% 60.3 17.6% 60.3 17.6% 60.3 17.6% 51.28 0.72 0.3612 60.8 7.8% 66.3 17.7% 66.3 17.7% 66.4 17.7% 56.37 1.85 0.9313 71.1 -4.8% 77.6 3.9% 87.6 17.3% 87.3 16.9% 74.69 0.63 0.3214 72.7 -0.3% 79.2 8.7% 87.5 20.0% 87.1 19.5% 72.92 0.77 0.3915 41.4 3.4% 45.2 13.0% 45.2 13.0% 45.1 12.7% 40.02 0.41 0.2016 49.6 5.3% 54.1 15.0% 54.1 15.0% 54.1 14.9% 47.07 1.49 0.7417 59.7 3.5% 65.2 12.9% 65.2 12.9% 65.2 12.9% 57.72 0.47 0.2418 65.9 1.4% 71.9 10.6% 71.9 10.6% 71.9 10.6% 65.03 2.01 1.0019 67.6 -10.0% 73.7 -2.0% 86.8 15.5% 86.5 15.0% 75.18 1.38 0.6920 38.7 -3.4% 42.3 5.5% 42.3 5.5% 42.2 5.3% 40.08 0.97 0.4821 35.9 2.4% 39.3 11.8% 39.3 11.8% 39.1 11.3% 35.12 0.09 0.0422 67.3 9.4% 73.5 19.4% 73.5 19.4% 73.6 19.5% 61.56 1.65 0.8223 45.9 3.7% 50.1 13.2% 50.1 13.2% 50.0 13.0% 44.25 1.17 0.5824 55.4 3.2% 60.4 12.6% 60.4 12.6% 60.4 12.6% 53.64 1.37 0.6925 50.3 2.3% 55.0 11.7% 55.0 11.7% 54.9 11.6% 49.20 0.37 0.1926 41.3 2.2% 45.1 11.6% 45.1 11.6% 45.0 11.4% 40.44 0.77 0.3827 61.0 2.0% 66.0 10.3% 66.0 10.3% 66.0 10.3% 59.82 0.54 0.2728 46.0 2.0% 50.2 11.5% 50.2 11.5% 50.2 11.4% 45.06 0.74 0.3729 46.5 4.6% 50.7 14.2% 50.7 14.2% 50.7 14.1% 44.44 1.28 0.6430 47.4 1.7% 51.8 11.0% 51.8 11.0% 51.7 10.9% 46.64 2.35 1.1731 66.0 1.6% 71.9 10.8% 71.9 10.8% 72.0 10.8% 64.93 2.55 1.2832 67.9 2.9% 74.1 12.2% 74.1 12.2% 74.1 12.2% 66.04 4.03 2.0233 36.2 1.9% 39.5 11.4% 39.5 11.4% 39.4 11.0% 35.49 0.28 0.1434 44.6 2.6% 48.7 12.1% 48.7 12.1% 48.7 11.9% 43.49 1.08 0.5435 53.9 1.0% 58.7 10.2% 58.7 10.2% 58.7 10.1% 53.32 1.35 0.6736 64.8 1.2% 70.7 10.3% 70.7 10.3% 70.7 10.3% 64.10 2.18 1.0937 64.7 -11.7% 70.4 -3.9% 86.0 17.4% 85.7 16.9% 73.29 0.27 0.1338 62.3 4.9% 68.0 14.4% 68.0 14.4% 68.0 14.5% 59.40 1.06 0.5339 57.1 -0.4% 62.3 8.7% 62.3 8.7% 62.3 8.7% 57.28 2.19 1.0940 63.2 4.7% 69.0 14.3% 69.0 14.3% 69.0 14.3% 60.35 0.36 0.1841 48.4 8.7% 52.8 18.7% 52.8 18.7% 52.7 18.6% 44.48 0.11 0.06

average 3.2% average 12.6% average 14.2% average 14.1%min -11.7% min -3.9% min 5.5% min 5.3%max 9.4% max 19.4% max 20.0% max 19.5%

TLDsCCEModel 0 Model 1 Model 2 Model 3

Table 42: The doses calculated using the CCE algorithm for each beam model, at each TLD position and a

comparison with the doses measured by the TLDs. For each model, the first column shows the calculated

dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the


97

5.4 Anthropomorphic phantom TLD (results corrected for revised TLD

calibrated in large phantom) comparison

dose stdev st uncert1 61.9 5.1% 66.3 12.6% 66.3 12.6% 66.3 12.6% 66.4 12.7% 58.9 2.23 1.112 66.4 6.0% 71.1 13.5% 71.1 13.5% 71.1 13.5% 72.1 15.1% 62.7 0.92 0.463 73.1 4.1% 78.1 11.2% 78.1 11.2% 78.1 11.2% 80.4 14.4% 70.3 3.71 1.854 60.2 7.5% 64.4 15.2% 64.4 15.2% 64.4 15.2% 64.5 15.2% 55.9 0.10 0.055 71.2 3.4% 76.1 10.6% 76.1 10.6% 76.1 10.6% 77.9 13.2% 68.9 2.42 1.216 73.7 2.7% 78.8 9.9% 79.4 10.8% 79.3 10.6% 81.0 12.9% 71.7 1.32 0.667 62.4 6.1% 66.8 13.6% 66.8 13.6% 66.8 13.6% 66.8 13.6% 58.8 0.95 0.488 37.8 5.0% 40.4 12.1% 40.4 12.1% 40.4 12.1% 37.8 5.0% 36.0 0.16 0.089 36.8 7.9% 39.3 15.2% 39.3 15.2% 39.3 15.2% 36.7 7.7% 34.1 0.50 0.2510 48.5 4.4% 51.9 11.7% 51.9 11.7% 51.9 11.7% 49.5 6.6% 46.5 0.54 0.2711 57.2 6.1% 61.1 13.5% 61.1 13.5% 61.1 13.5% 59.5 10.5% 53.8 0.72 0.3612 63.9 7.9% 68.3 15.5% 68.3 15.5% 68.3 15.5% 68.1 15.0% 59.2 1.85 0.9313 73.2 -6.6% 78.5 0.0% 88.6 12.9% 88.2 12.4% 88.7 13.2% 78.4 0.63 0.3214 74.9 -2.1% 80.3 4.8% 88.5 15.6% 88.1 15.1% 89.0 16.2% 76.6 0.77 0.3915 42.6 1.3% 45.5 8.2% 45.5 8.2% 45.5 8.2% 42.8 1.8% 42.0 0.41 0.2016 51.3 3.8% 54.9 11.0% 54.9 11.0% 54.9 11.0% 52.9 7.0% 49.4 1.49 0.7417 60.4 -0.4% 64.6 6.6% 64.6 6.6% 64.6 6.6% 63.9 5.4% 60.6 0.47 0.2418 68.5 0.3% 73.3 7.3% 73.3 7.3% 73.3 7.3% 74.5 9.1% 68.3 2.01 1.0019 68.3 -13.5% 73.1 -7.4% 86.2 9.3% 85.9 8.8% 86.1 9.1% 78.9 1.38 0.6920 42.9 2.0% 45.9 9.2% 45.9 9.2% 45.9 9.2% 43.6 3.6% 42.1 0.97 0.4821 39.5 7.1% 42.2 14.4% 42.2 14.4% 42.2 14.4% 39.6 7.4% 36.9 0.09 0.0422 71.0 9.9% 76.0 17.7% 76.0 17.7% 76.0 17.7% 76.9 19.0% 64.6 1.65 0.8223 47.8 2.9% 51.2 10.2% 51.2 10.2% 51.2 10.2% 49.8 7.3% 46.5 1.17 0.5824 57.3 1.8% 61.4 9.1% 61.4 9.1% 61.4 9.1% 61.0 8.3% 56.3 1.37 0.6925 52.2 1.1% 56.0 8.4% 56.0 8.4% 56.0 8.4% 55.0 6.5% 51.7 0.37 0.1926 43.4 2.2% 46.5 9.4% 46.5 9.4% 46.5 9.4% 44.7 5.2% 42.5 0.77 0.3827 63.0 0.3% 67.0 6.7% 67.0 6.7% 67.0 6.7% 67.0 6.7% 62.8 0.54 0.2728 47.6 0.6% 50.9 7.6% 50.9 7.6% 50.9 7.6% 48.5 2.4% 47.3 0.74 0.3729 47.9 2.7% 51.3 9.9% 51.3 9.9% 51.3 9.9% 48.9 4.9% 46.7 1.28 0.6430 49.4 1.0% 52.9 8.0% 52.9 8.0% 52.9 8.0% 50.8 3.7% 49.0 2.35 1.1731 68.2 0.1% 73.0 7.1% 73.0 7.1% 73.0 7.1% 74.1 8.7% 68.2 2.55 1.2832 70.3 1.4% 75.2 8.5% 75.2 8.5% 75.2 8.5% 76.5 10.3% 69.3 4.03 2.0233 38.3 2.7% 41.0 10.0% 41.0 10.0% 41.0 10.0% 38.9 4.4% 37.3 0.28 0.1434 46.6 2.0% 49.9 9.2% 49.9 9.2% 49.9 9.2% 48.2 5.5% 45.7 1.08 0.5435 55.9 -0.2% 59.8 6.9% 59.8 6.9% 59.8 6.9% 58.9 5.3% 56.0 1.35 0.6736 66.2 -1.6% 70.9 5.3% 70.9 5.3% 70.9 5.3% 71.3 5.9% 67.3 2.18 1.0937 63.5 -17.4% 67.9 -11.7% 83.4 8.4% 83.1 8.0% 82.3 7.0% 77.0 0.27 0.1338 62.5 0.2% 66.9 7.2% 66.9 7.2% 66.9 7.2% 66.8 7.2% 62.4 1.06 0.5339 58.1 -3.3% 62.2 3.4% 62.2 3.4% 62.2 3.4% 60.9 1.3% 60.1 2.19 1.0940 66.1 4.3% 70.8 11.7% 70.8 11.7% 70.8 11.7% 71.2 12.4% 63.4 0.36 0.1841 50.9 9.0% 54.4 16.6% 54.4 16.6% 54.4 16.6% 52.4 12.2% 46.7 0.11 0.06

average 1.9% average 9.0% average 10.5% average 10.5% average 8.7%min -17.4% min -11.7% min 3.4% min 3.4% min 1.3%max 9.9% max 17.7% max 17.7% max 17.7% max 19.0%

TLDsPBEModel 0 Model 1 Model 2 Model 3 Model 4

Table 43: The doses calculated using the pencil beam algorithm for all models and the percentage differences

between the measured and calculated values. For each model, the first column shows the calculated dose in

cGy for 1000 MU, and the second column shows the percentage deviation of that dose from the revised


98

dose stdev st uncert1 59.9 1.6% 65.3 10.8% 65.3 10.8% 65.3 10.9% 58.9 2.23 1.112 64.7 3.2% 70.5 12.6% 70.5 12.6% 70.6 12.6% 62.7 0.92 0.463 72.4 3.0% 78.7 12.0% 78.7 12.0% 78.8 12.1% 70.3 3.71 1.854 57.4 2.5% 62.6 11.9% 62.6 11.9% 62.6 11.9% 55.9 0.10 0.055 69.8 1.4% 76.1 10.5% 76.1 10.5% 76.1 10.6% 68.9 2.42 1.216 73.4 2.3% 80.1 11.7% 80.7 12.5% 80.7 12.5% 71.7 1.32 0.667 60.7 3.2% 66.3 12.7% 66.3 12.7% 66.3 12.7% 58.8 0.95 0.488 36.0 -0.1% 39.3 9.2% 39.3 9.2% 39.2 8.8% 36.0 0.16 0.089 35.2 3.3% 38.5 12.9% 38.5 12.9% 38.3 12.5% 34.1 0.50 0.25

10 46.5 0.0% 50.7 9.2% 50.7 9.2% 50.7 9.0% 46.5 0.54 0.2711 55.2 2.6% 60.3 12.0% 60.3 12.0% 60.3 12.0% 53.8 0.72 0.3612 60.8 2.7% 66.3 12.1% 66.3 12.1% 66.4 12.1% 59.2 1.85 0.9313 71.1 -9.3% 77.6 -1.1% 87.6 11.8% 87.3 11.3% 78.4 0.63 0.3214 72.7 -5.1% 79.2 3.5% 87.5 14.3% 87.1 13.8% 76.6 0.77 0.3915 41.4 -1.6% 45.2 7.6% 45.2 7.6% 45.1 7.4% 42.0 0.41 0.2016 49.6 0.3% 54.1 9.5% 54.1 9.5% 54.1 9.4% 49.4 1.49 0.7417 59.7 -1.5% 65.2 7.5% 65.2 7.5% 65.2 7.5% 60.6 0.47 0.2418 65.9 -3.4% 71.9 5.3% 71.9 5.3% 71.9 5.4% 68.3 2.01 1.0019 67.6 -14.3% 73.7 -6.6% 86.8 10.0% 86.5 9.6% 78.9 1.38 0.6920 38.7 -8.0% 42.3 0.5% 42.3 0.5% 42.2 0.3% 42.1 0.97 0.4821 35.9 -2.5% 39.3 6.5% 39.3 6.5% 39.1 6.0% 36.9 0.09 0.0422 67.3 4.2% 73.5 13.7% 73.5 13.7% 73.6 13.8% 64.6 1.65 0.8223 45.9 -1.3% 50.1 7.8% 50.1 7.8% 50.0 7.6% 46.5 1.17 0.5824 55.4 -1.7% 60.4 7.2% 60.4 7.2% 60.4 7.2% 56.3 1.37 0.6925 50.3 -2.6% 55.0 6.4% 55.0 6.4% 54.9 6.3% 51.7 0.37 0.1926 41.3 -2.6% 45.1 6.3% 45.1 6.3% 45.0 6.1% 42.5 0.77 0.3827 63.0 0.3% 67.0 6.7% 67.0 6.7% 67.0 6.7% 62.8 0.54 0.2728 46.0 -2.8% 50.2 6.2% 50.2 6.2% 50.2 6.1% 47.3 0.74 0.3729 46.5 -0.4% 50.7 8.8% 50.7 8.8% 50.7 8.6% 46.7 1.28 0.6430 47.4 -3.1% 51.8 5.7% 51.8 5.7% 51.7 5.6% 49.0 2.35 1.1731 66.0 -3.3% 71.9 5.5% 71.9 5.5% 72.0 5.6% 68.2 2.55 1.2832 67.9 -2.0% 74.1 6.8% 74.1 6.8% 74.1 6.9% 69.3 4.03 2.0233 36.2 -2.9% 39.5 6.1% 39.5 6.1% 39.4 5.7% 37.3 0.28 0.1434 44.6 -2.3% 48.7 6.7% 48.7 6.7% 48.7 6.6% 45.7 1.08 0.5435 53.9 -3.8% 58.7 4.9% 58.7 4.9% 58.7 4.9% 56.0 1.35 0.6736 64.8 -3.7% 70.7 5.0% 70.7 5.0% 70.7 5.1% 67.3 2.18 1.0937 64.7 -15.9% 70.4 -8.5% 86.0 11.8% 85.7 11.4% 77.0 0.27 0.1338 62.3 -0.1% 68.0 8.9% 68.0 8.9% 68.0 9.0% 62.4 1.06 0.5339 57.1 -5.1% 62.3 3.5% 62.3 3.5% 62.3 3.5% 60.1 2.19 1.0940 63.2 -0.2% 69.0 8.8% 69.0 8.8% 69.0 8.9% 63.4 0.36 0.1841 48.4 3.6% 52.8 13.1% 52.8 13.1% 52.7 13.0% 46.7 0.11 0.06

average -1.6% average 7.3% average 8.8% average 8.7%min -15.9% min -8.5% min 0.5% min 0.3%max 4.2% max 13.7% max 14.3% max 13.8%

TLDsCCEModel 0 Model 1 Model 2 Model 3

Table 44: The doses calculated using the Collapsed Cone algorithm for all models and the percentage

differences between the measured and calculated values. For each model, the first column shows the

calculated dose in cGy for 1000 MU, and the second column shows the percentage deviation of that dose

from the revised TLD-measured dose at that point for the same number of MU.

99

REFERENCES

1. Brown, J., et al., Increasing Incidence of Late Second Malignancies After Conditioning With Cyclophosphamide and Total-Body Irradiation and Autologous Bone Marrow Transplantation for Non-Hodgkin's Lymphoma. Journal of Clinical Oncology, 2005. 23(10): p. 2208-2214.

2. De Neve, W., M. Lybeert, and J. Meerwaldt, Low-Dose Total Body Irradiation in Non-Hodgkin Lymphoma: Short-and Long-term Toxicity and Prognostic Factors. American Journal of Clinical Oncology, 1990. 13(4): p. 280-284.

3. Mukai, H.Y., et al., Successful treatment of a patient with subcutaneous panniculitis-like T-cell lymphoma with high-dose chemotherapy and total body irradiation. European Journal of Haematology, 2003. 70(6): p. 413-416.

4. Clift, R.A., et al., Allogeneic Marrow Transplantation in Patients With Acute Myeloid Leukemia in First Remission: A Randomized Trial of Two Irradiation Regimens. Blood, 1990. 76: p. 1867-1871.

5. Clift, R.A., et al., Allogeneic Marrow Transplantation in Patients With Chronic Myeloid Leukemia in the Chronic Phase: A Randomized Trial of Two Irradiation Regimens. Blood, 1991. 77(8): p. 1660-1665.

6. Alexander, B.M., et al., Utility Of Cranial Boost In Addition To Total Body Irradiation In The Treatment Of High Risk Acute Lymphoblastic Leukemia. International Journal of Radiation Oncology Biology Physics, 2005. 63(4): p. 1191-1196.

7. Inagaki, J., et al., Bone marrow transplantation in children with severe aplastic anemia using a conditioning regimen containing 3 Gy of total body irradiation, cyclophosphamide with or without antithymocyte globulin. Pediatric Transplantation, 2007. 11(2): p. 180-186.

8. Van Dyk, J., et al., The physical aspects of total and half body photon irradiation: a report of Task Group 29, Radiation Therapy Committee Association of Physicists in Medicine. 1986: American Association of of Physicists in Medicine.

9. Harden, S.V., et al., Total body irradiation using a modified standing technique: a single institution 7 year experience. British Journal of Radiology, 2001. 74(887): p. 1041-1047.

10. Maynes, P., A. Nahum, and J.-C. Rosenwald, Handbook of Radiotherapy Physics: Theory and Practice. 2007, Boca Raton, FL, US: Taylor and Francis Publishing Group

11. Shank, B., Total Body Irradiation for Marrow or Stem-Cell Transplantation. Cancer Investigation, 1998. 16(6): p. 397 - 404.

12. Metcalfe, P., T. Kron, and P. Hoban, The physics of radiotherapy x-rays and electrons. 2007, Madison: Medical Physics Publishing. xix, 905 p. :.

13. Khan, F.M. and e. al, Treatment Planning in Radiation Oncology. Second ed, ed. F.M. Khan. 2007, Philadelphia, PA, USA: Lippincott Williams and Wilkins.

14. Van Dyk, J., The Modern Technology of Radiation Oncology. Vol. 1. 1999, Madison, Wisconsin, USA: Medical Physics Publishing.

100

15. Millar, M., et al., ACPSEM position paper recommendations for the safe use of external beams and sealed brachytherapy sources in radiation oncology. Australasian Physical and Engineering Sciences in Medicine, 1997. 20.

16. Mayles, W., et al., Physics aspects of quality control in radiotherapy. 1999: Institute of Physics and Engineering in Medicine.

17. Hui, S.K., et al., CT-based analysis of dose homogeneity in total body irradiation using lateral beam. Journal of Applied Clinical Medical Physics, 2004. 5(4): p. 71-79.

18. Lagrange, J.L., et al., CT Measurement of Lung Density: The Role of Patient Position and Value for Total Body Irradiation. International Journal of Radiation Oncology Biology Physics, 1987. 13(5): p. 941-944.

19. Hurkmans, C., et al., Limitations of a pencil beam approach to photon dose calculations in the head and neck region. Radiotherapy and Oncology, 1995. 37(1): p. 74-80.

20. Knöös, T., et al., Limitations of a pencil beam approach to photon dose calculations in lung tissue. Physics in Medicine and Biology, 1995. 40(9): p. 1411-1420.

21. Wieslander, E. and T. Knöös, Dose perturbation in the presence of metallic implants: treatment planning system versus Monte Carlo simulations. Physics in Medicine and Biology, 2003. 48(20): p. 3295-3305.

22. Knöös, T., et al., Comparison of dose calculation algorithms for treatment planning in external photon beam therapy for clinical situations. Physics in Medicine and Biology, 2006. 51(22): p. 5785-5807.

23. Mackie, T.R., et al., Generation of photon energy deposition kernels using the EGS Monte Carlo code. Physics in Medicine and Biology, 1988. 33(1): p. 1-20.

24. Wieslander, E. and T. Knöös, Monte Carlo based verification of a beam model used in a treatment planning system. Journal of Physics: Conference Series, 2008. 102: p. 012027.

25. Knöös, T., et al., The dosimetric verification of a pencil beam based treatment planning system. Physics in Medicine and Biology, 1994. 39(10): p. 1609-1628.

26. Aspradakis, M.M., et al., Experimental verification of convolution/superposition photon dose calculations for radiotherapy treatment planning. Physics in Medicine and Biology, 2003. 48(17): p. 2873-2893.

27. Aspradakis, M.M., H.M. McCallum, and N. Wilson, Dosimetric and treatment planning considerations for radiotherapy of the chest wall. The British Journal of Radiology, 2006. 79(946): p. 828-836.

28. Ahnesjö, A. and M.M. Aspradakis, Dose calculations for external photon beams in radiotherapy. Physics in Medicine and Biology, 1999. 44(11): p. R99-R155.

29. Fotina, I., et al., Advanced kernel methods vs. Monte Carlo-based dose calculation for high energy photon beams. Radiotherapy and Oncology, 2009. 93(3): p. 645-653.

30. Paelinck, L., et al., Experimental verification of lung dose with radiochromic film: comparison with Monte Carlo simulations and commercially available treatment planning systems. Physics in Medicine and Biology, 2005. 50(9): p. 2055-2069.

31. Irvine, C., et al., The Clinical Implications of the Collapsed Cone Planning Algorithm. Clinical Oncology, 2004. 16(2): p. 148-154.

32. Ahnesjö, A. and A. Trepp, Acquisition of the effective lateral energy fluence distribution for photon beam dose calculations by convolution models. Physics in Medicine and Biology, 1991. 36(7): p. 973.

33. Ahnesjö, A., M. Saxner, and A. Trepp, A pencil beam model for photon dose calculation. Medical Physics, 1992. 19(2): p. 263-273.

101

34. Ahnesjö, A., et al., Beam modeling and verification of a photon beam multisource model. Medical Physics, 2005. 32(6): p. 1722-1737.

35. Ahnesjö, A., Collapsed cone convolution of radiant energy for photon dose calculation in heterogeneous media. Medical Physics, 1989. 16(4): p. 577-592.

36. Khan, F.M., The Physics of Radiation Therapy. 3rd ed. 2003, Philadelphia, PA: Lippincott Williams & Wilkins.

37. Van Dyk, J., The Modern Technology of Radiation Oncology. Vol. 2. 2005, Madison: Medical Physics Publishing.

38. Knoll, G.F., Radiation Detection and Measurement. Third ed. 2000, USA: John Wiley & Sons.

39. Podgorsak, E.B. and e. al, Radiation Oncology Physics: A Handbook for Teachers And Students. 2005: International Atomic Energy Agency.

40. Duch, M.A., et al., Thermoluminescence dosimetry applied to in vivo dose measurements for total body irradiation techniques. Radiotherapy and Oncology, 1998. 47(3): p. 319-324.

41. Maher, J.C.P.H.N.A.L.K.P., Applied Imaging Technology. Fourth ed. 2001, Melbourne, Australia.

42. Deeg, H.J., et al., Single Dose or Fractionated Total Body Irradiation and Autologous Marrow Transplantation in Dogs: Effects of Exposure Rate, Fraction Size, and Fractionation Interval on Acute and Delayed Toxicity. International Journal of Radiation Oncology Biology Physics, 1988. 15(3): p. 647-653.

43. Cosset, J.M., et al., Single dose versus hyperfractionated total body irradiation before allogeneic bone marrow transplantation: a non-randomized comparative study of 54 patients at the Institut Gustave-Roussy. Radiotherapy and Oncology, 1989. 15(2): p. 151-160.

44. Cosset, J.-M., et al., Single Dose Versus Fractionated Total Body Irradiation Before Bone Marrow Transplantation: Radiobiological and Clinical Considerations. International Journal of Radiation Oncology Biology Physics, 1994. 30(2): p. 477-492.

45. Cox, J.D. and K.K. Ang, Radiation Oncology: Rationale, Techniques, Results. Eighth ed. 2003, St Louis, USA: Mosby Inc.

46. Keane, T., J. Van Dyk, and W. Rider, Idiopathic interstitial pneumonia following bone marrow transplantation: the relationship with total body irradiation. International Journal of Radiation Oncology* Biology* Physics, 1981. 7(10): p. 1365-1370.

47. Van Dyk, J., et al., Radiation Pneumonitis Following Large Single Dose Irradiation: A Re-evaluation Based On Absolute Dose To Lung. International Journal of Radiation Oncology Biology Physics, 1981. 7(4): p. 461-467.

48. O'Donoghue, J.A., Fractionated versus low dose-rate total body irradiation.Radiobiological considerations in the selection of regimes. Radiotherapy and Oncology, 1986. 7(3): p. 241-247.

49. ICRU, ICRU Report 24: Determination of Absorbed Dose in a Patient Irradiated by Beams of X or Gamma Rays in Radiotherapy Procedures. 1976, International Commission on Radiation Units and Measurements (ICRU).

50. ICRU, ICRU Report 50: Prescribing, Recording and Reporting Photon Beam Therapy. 1993, International Commission on Radiation Units and Measurements (ICRU).

51. Safwat, A., et al., The clinical, physical and radiobiological aspects of total body irradiation. European Journal of Cancer, 1995. 31(Supplement 6): p. S33-S33.

52. Humphries, S.M., et al., Comparison of super stuff and paraffin wax bolus in radiation therapy of irregular surfaces. Medical Dosimetry, 1996. 21(3): p. 155-157.

102

53. Pla, M., S.G. Chenery, and E.B. Podgorsak, Total body irradiation with a sweeping beam. International Journal of Radiation Oncology*Biology*Physics, 1983. 9(1): p. 83-89.

54. Quast, U., Total body irradiation--review of treatment techniquesin Europe. Radiotherapy and Oncology, 1987. 9(2): p. 91-106.

55. Breneman, J.C., et al., A technique for delivery of total body irradiation for bone marrow transplantation in adults and adolescents. International Journal of Radiation Oncology*Biology*Physics, 1990. 18(5): p. 1233-1236.

56. Niroomand-Rad, A., Physical aspects of total body irradiation of bone marrow transplant patients using 18 MV X rays. International Journal of Radiation Oncology*Biology*Physics, 1991. 20(3): p. 605-611.

57. Miralbell, R., et al., Can a total body irradiation technique be fast and reproducible? International Journal of Radiation Oncology*Biology*Physics, 1994. 29(5): p. 1167-1173.

58. Jones, D., et al., An isocentrically mounted stand for total body irradiation. British Journal of Radiology, 2000. 73(871): p. 776-779.

59. Umek, B., M. Zwitter, and H. Habic, Total body irradiation with translation method. Radiotherapy and Oncology, 1996. 38(3): p. 253-255.

60. Gerrard, G.E., et al., Toxicity and dosimetry of fractionated total body irradiation prior to allogeneic bone marrow transplantation using a straightforward radiotherapy technique. Clinical Oncology, 1998. 10(6): p. 379-383.

61. Sampath, S., T.E. Schultheiss, and J. Wong, Dose response and factors related to interstitial pneumonitis after bone marrow transplant. International Journal of Radiation Oncology*Biology*Physics, 2005. 63(3): p. 876-884.

62. Gerbi, B.J. and K.E. Dusenbery, Design specifications for a treatment stand used for total body photon irradiation with patients in a standing position. Medical Dosimetry, 1995. 20(1): p. 25-30.

63. Papiez, E., et al., A Novel Technique for Total Body Irradiation using Gantry Rotation and Beam Modulation. International Journal of Radiation Oncology*Biology*Physics, 2008. 72(1, Supplement 1): p. S677-S677.

64. La Macchia, N.A., et al., The Development of a Segmented Intensity Modulated Beam Technique for Total Body Irradiation Prior to Hematologic Stem Cell Transplantation. International Journal of Radiation Oncology*Biology*Physics, 2007. 69(3, Supplement 1): p. S540-S540.

65. Quast, U., Physical treatment planning of total-body irradiation: Patient translation and beam-zone method. Medical Physics, 1985. 12(5): p. 567-574.

66. Harden, S.V., et al., Total body irradiation using a modified standing technique: a single institution 7 year experience. British Journal of Radiology, 2001. 74: p. 1041- 1047.

67. Sánchez-Nieto, B., F. Sánchez-Doblado, and J.A. Terrón, A CT-aided PC-based physical treatment planning of TBI: a method for dose calculation. Radiotherapy and Oncology, 1997. 42(1): p. 77-85.

68. Houdek, P.V. and V.J. Pisciotta, A comparison of calculated and measured data for total body irradiation by 10 MV x-rays. Physics in Medicine and Biology, 1987. 32(9): p. 1101-1108.

69. Curran, W.J., J.M. Galvin, and G.J. D'Angio, A simple dose calculation method for total body photon irradiation. International Journal of Radiation Oncology*Biology*Physics, 1989. 17(1): p. 219-224.

103

70. Podgorsak, E.B., et al., The influence of phantom size on output, peak scatter factor, and percentage depth dose in large-field photon irradiation. Medical Physics, 1985. 12(5): p. 639-645.

71. Low, D.A., et al., A technique for the quantitative evaluation of dose distributions. Medical Physics, 1998. 25(5): p. 656-661.

72. Andreo, P., et al., TRS 398: Absorbed Dose Determination in External Beam Radiotherapy: An International Code of Practice for Dosimetry based on Standards of Absorbed Dose to Water. 2004, International Atomic Energy Agency (IAEA): Vienna, Austria.

73. Andreo, P., et al., TRS 430: Commissioning and quality assurance of computerized planning systems for radiation treatment of cancer. 2004, International Atomic Energy Agency (IAEA): Vienna, Austria.

74. Poole, C., Using Monte Carlo simulation to predict dose at extended source to surface distance, in School of Physical and Chemical Sciences. 2008, Queensland University of Technology: Brisbane.

2012 commissioning of a new total body irradiation protocol

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