advances in very small x ray field dosimetry for … · mm circular fields for which limited...

ADVANCES IN VERY SMALL X-RAY FIELD

DOSIMETRY FOR CIRCULAR CONES USED

IN STEREOTACTIC RADIOSURGERY

Johnny Estuardo Morales

(MSc)

Dr Jamie Trapp and Dr Scott Crowe

Submitted in fulfilment of the requirements for the degree of

Doctor of Philosophy

School of Chemistry, Physics and Engineering

Science and Engineering Faculty

Queensland University of Technology

2019

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery i

Keywords

Medical physics, Monte Carlo, small field dosimetry, radiation oncology,

stereotactic radiosurgery, microDiamond, EBT3, Gafchromic film, Brainlab, iPlan,

Brainlab cones, OSLDs, IAEA TRS-483

ii Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

Abstract

Advances in new technology in radiation oncology require innovation to

implement new methods and new equipment for validation of very small field

dosimetry in stereotactic radiosurgery (SRS). Small field dosimetry presents a major

challenge due to issues relating to source occlusion, volume averaging and

perturbations caused in the field from the detectors used for these measurements.

The IAEA TRS-483 Code of Practice titled the Dosimetry of Small Static Fields

Used in External Beam Radiotherapy Code of Practice was published in 2017 and is

intended to serve as a guide for performing measurements for small x-ray fields.

However, the advice and recommendations of this Code of Practice was limited to

fields with dimensions of 5 mm or greater. This is consistent with the literature in that

most publications have concentrated their effort in fields greater than 5 mm However,

there are situations in clinical practice where smaller field sizes are used for treatment.

For example, the use of a 4-mm diameter conical applicator is an effective tool in the

treatment of Trigeminal Neuralgia. Such treatments usually involved the delivery of

90 Gy in a single fraction.

The objective of this thesis was to provide dosimetry advice and develop new

techniques to be used with new and existing detectors for small fields, including the 4-

mm circular fields for which limited recommendations were available. These

objectives were achieved through 5 publications.

The suitability of a new synthetic diamond detector, the PTW 60019

microDiamond for dosimetry in circular defined fields of 4 to 30 mm in diameter was

established. Field output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were derived at a depth of 1.5

cm in water at a Source to Surface Distance (SSD) of 100 cm using a 6 MV SRS x-ray

beam on a Novalis Tx linear accelerator. At the time of publication of this work, there

were no published correction factors for this detector. The field correction factors were

determined to be within 2.7% which is consistent with current guidelines.

Field output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for Gafchromic EBT3 radiochromic

film in water were determined through Monte Carlo simulations using the 6 MV SRS

x-ray beam, which was modelled using BEAMnrc/DOSXYZnrc software code. Most

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery iii

publications have worked on the assumption that EBT3 film is a radiologically water

equivalent detector, however there had not been any Monte Carlo simulations have

been published to show that EBT3 film is radiologically water equivalent for small x-

ray fields until this work. The results showed that the field correction factors for

Gafchromic EBT3 film were less than 1% over all the field sizes studied down to 4

mm diameter.

A novel extrapolation technique was tested and published on the use of

Gafchromic EBT3 film measurements to eliminate the volume averaging effect for

small fields of 4 mm diameter. At such a small field size, the volume averaging effect

can be a major issue in the dosimetry as has been extensively demonstrated in the

literature. This technique allowed the estimation of the output factor for a 4 mm

Brainlab circular cone in water. The value obtained using this technique was 0.649.

The extrapolation technique published and presented in this thesis using

Gafchromic EBT3 film was also implemented in another promising detector for small

field dosimetry, the Optically Stimulated Luminescence detector (OSLDs). The

original OSLD detectors were modified to reduce their effective area so as to also

eliminate the volume averaging effects. The work in this thesis was an improvement

of previous work in this area which had only performed their work with a minimum of

field size of 7.5 mm. The work in this thesis examined field sizes of 4 mm diameter as

produced by the Brainlab cones and used in the treatment of trigeminal neuralgia

nerves in patients.

Relative skin or surface doses for the Brainlab circular cones were characterised

through measurement. To date, there is limited data available in the literature in terms

of the skin dose for the very small field sizes used in SRS treatments. It is important

to note that current commercially available treatment planning systems cannot account

for or predict dose delivered to the skin at depths near the surface. The work in this

thesis fills this gap in knowledge by providing the surface dose values as percentage

of the dose at the depth of dose maxim for x-ray fields used in SRS treatments using

the Brainlab cones. The skin dose values, as defined by the ICRU to be 70 m deep,

were between 13 – 15 % of the value at dmax for the circular cones from 4 to 30 mm in

diameter.

iv Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

Finally, the accuracy of commercial treatment planning system algorithm (iPlan)

for use with Brainlab circular cones for inhomogenities present in phantom slabs was

evaluated. There is limited published literature on the effect of the iPlan algorithm for

circular cones. Most published literature has concentrated on the use of multi-leaf

collimators rather than circular cones, despite the fact that some of the highest single

fraction doses are delivered with circular cones. A study was performed to evaluate

the performance of this planning system in an anthropomorphic phantom to simulate

a patient treatment in the presence on a low density cavity region.

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery v

Table of Contents

Keywords .................................................................................................................................. i

Abstract .................................................................................................................................... ii

Table of Contents ......................................................................................................................v

List of Figures ....................................................................................................................... viii

List of Tables .......................................................................................................................... xi

List of Publications for PhD Candidature .............................................................................. xii

List of Other Research Activities and Published Works during PhD Candidature ............... xiii

Statement of Original Authorship ......................................................................................... xiv

Acknowledgements .................................................................................................................xv

Chapter 1: Introduction ...................................................................................... 1

1.1 Radiation therapy and stereotactic radiosurgery .............................................................1

1.2 Dosimetry in radiotherapy ..............................................................................................5

1.3 Literature review – What have we learnt? ......................................................................8

1.4 Aims of the project .......................................................................................................12

1.5 Thesis outline ................................................................................................................13

1.6 References ....................................................................................................................16

Statement of Co-Authors for Chapter 2 ................................................................ 23

Chapter 2: Dosimetry of cone-defined stereotactic radiosurgery fields with a

commercial synthetic diamond detector ................................................................ 25

2.1 Introduction ..................................................................................................................28

2.2 Methods and materials ..................................................................................................29

2.3 Results ..........................................................................................................................30

2.4 Discussion .....................................................................................................................35

2.5 Conclusion ....................................................................................................................37

2.6 Acknowledgements.......................................................................................................37

2.7 Conflict of interest ........................................................................................................37

2.8 References ....................................................................................................................37

Statement of Co-Authors for Chapter 3 ................................................................ 41

Chapter 3: Monte Carlo calculated output correction factors for

Gafchromic EBT3 film for dosimetry in stereotactic radiosurgery ray fields ... 43

3.1 Introduction ..................................................................................................................46

3.2 Materials and methods ..................................................................................................48

3.3 Results ..........................................................................................................................49

3.4 Discussion .....................................................................................................................50

vi Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

3.5 Conclusion ................................................................................................................... 52

3.6 Acknowledgements ...................................................................................................... 52

3.7 Conflict of interest ....................................................................................................... 53

3.8 References .................................................................................................................... 53

Statement of Co-Authors for Chapter 4 ................................................................. 55

Chapter 4: An experimental extrapolation technique using Gafchromic

EBT3 film for relative output factor measurements in small x-ray fields .......... 57

4.1 Introduction .................................................................................................................. 60

4.2 Materials and methods ................................................................................................. 60

4.3 Results .......................................................................................................................... 62

4.4 Discussion .................................................................................................................... 68

4.5 Conclusions .................................................................................................................. 70

4.6 Conflict of Interest ....................................................................................................... 70

4.7 References .................................................................................................................... 70


Chapter 5: A novel extrapolation method using OSL detectors for very small

field output factor measurement for stereotactic radiosurgery ........................... 77

5.1 Introduction .................................................................................................................. 80

5.2 Materials and Methods ................................................................................................. 81

5.3 Results .......................................................................................................................... 84

5.4 Discussion .................................................................................................................... 87

5.5 Conclusion ................................................................................................................... 88

5.6 Conflict of interest ....................................................................................................... 88

5.7 References .................................................................................................................... 89


Chapter 6: A comparison of surface doses for very small field x-ray beams:

Monte Carlo calculations and radiochromic film ................................................. 95

6.1 Introduction .................................................................................................................. 98

6.2 Materials and methods ................................................................................................. 99

6.3 Results and Discussion ............................................................................................... 102

6.4 Conclusions ................................................................................................................ 107

6.5 Acknowledgements .................................................................................................... 107

6.6 Conflict of interest ..................................................................................................... 107

6.7 References .................................................................................................................. 107

Statement of Co-Authors for Chapter 7 ............................................................... 111

Chapter 7: A study of dose inhomogeneity correction in a commercial

treatment planning system for stereotactic radiosurgery .................................. 113

7.1 Introduction ................................................................................................................ 116

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery vii

7.2 Materials and methods ................................................................................................118

7.3 Results ........................................................................................................................123

7.4 Discussion ...................................................................................................................128

7.5 Conclusion ..................................................................................................................130

7.6 Acknowledgements.....................................................................................................130

7.7 Compliance with Ethical Standards ............................................................................131

7.8 Conflict of interest ......................................................................................................131

7.9 References ..................................................................................................................131

Chapter 8: Discussions and Conclusions ....................................................... 135

8.1 Discussion on new commercial detector – PTW 60019 microdiamond .....................135

8.2 Discussion on Monte Carlo modelling of Gafchromic EBT3 film .............................137

8.3 Discussion on extrapolation technique for Gafchromic EBT3 film ...........................138

8.4 Discussion on extrapolation technique for OSLD detectors .......................................138

8.5 Discussion on skin/Surface dose for Brainlab circular cones .....................................139

8.6 Discussion on inhomogenenity correction on small fields produced by Brainlab

circular cones and implementation of MMCTP for monte carlo based independent checks 139

8.7 Clinical implications of the work................................................................................140

8.8 Conclusions ................................................................................................................141

8.9 Future work in small field dosimetry for radiosurgery ...............................................142

8.10 References ..................................................................................................................143

viii Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

List of Figures

Figure 1-1 Novalis Tx linear accelerator capable of producing 6 MV x-rays for

treatment in radiotherapy ............................................................................... 3

Figure 1-2 A set of Brainlab circular cones used for SRS treatments on a linear

accelerator. ..................................................................................................... 4

Figure 2-1 Percentage depth doses measured with a PTW 60019

microDiamond detector (O), IBA SFD Diode (X) and PTW 60012 E

Diode detector (+) for 4, 7.5, 10 and 30 mm circular cones at SSD 100

cm. ................................................................................................................ 31

Figure 2-2 Half cross profiles measured at a depth of 10.0 cm with a PTW

60019 microDiamond detector in parallel orientation (),

perpendicular orientation (), and IBA SFD diode (X) for 4, 7.5, 10

and 30 mm circular cone at SSD of 100 cm. ............................................... 33

Figure 4-1 Measured and extrapolated output factor values for a 4mm Brainlab

cone using Gafchromic EBT3 film with varying sizes of analysis area.

As the area of analysis decreases, an increase in measured output

factor occurs due to the non-plateauing nature of the 4mm cone

profile. .......................................................................................................... 64

Figure 4-2 Measured net optical density profile for a 6MV SRS x-ray beam, 4

mm cone taken as the average of nine EBT3 film measurements. The

insert in the figure includes details of the centre 1 mm of the profile. ........ 64

Figure 4-3 Measured and extrapolated output factor values for a 25 mm

diameter Brainlab cone using Gafchromic EBT3 film with varying

diameters of analysis area. As the area of analysis decreases negligible

differences in measured output factor occurs due to the plateauing

nature of the 25mm cone profile. ................................................................. 65

Figure 4-4 Measured net optical density profile for a 6MV x-ray beam, 25 mm

diameter Brainlab cone. Insert in the picture includes details of the

centre 4 mm of the profile for an example film showing the central

plataeu effect at this field size. ..................................................................... 65

Figure 4-5 Percentage difference in output factor from the extrapolated zero

area value for the 4 and 25 mm Brainlab cones. Variations are seen

using the 4 mm cone but negligible differences calculated at larger

cone sizes. .................................................................................................... 66

Figure 4-6 Exaggerated example of a small field. The average value of pixels

in ROI A will be greater than that of ROI B. ............................................... 68

Figure 5-1 A demonstration photo of the modified OSLD’s (middle and right)

comparing with a standard nanoDot OSLD (left). ....................................... 82

Figure 5-2 A dosimetric the set up for modified OSLD’s with the tray slot

filled with liquid water. ................................................................................ 84

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery ix

Figure 5-3 Measured and extrapolated output factor values for a 6 MV SRS x-

ray beam with the 30 mm diameter Brainlab cone using the modified

OSLDs with varying hole sizes.................................................................... 85


ray beam with 10 mm Brainlab cone using the modified OSLDs with

varying hole sizes. ........................................................................................ 85


ray beam with the 7.5 mm Brainlab cone using the modified OSLDs

with varying hole sizes................................................................................. 86


ray beam with the 4 mm Brainlab cone using the modified OSLDs

with varying hole sizes................................................................................. 86

Figure 6-1 Percentage depth dose in water calculated for a 10 × 10 cm2 field.

BEAMnrc/DOSXYZnrc versus measurement by an Advanced Markus

ionisation chamber. .................................................................................... 103

Figure 6-2 a) Cross profiles at depth in water calculated by

BEAMnrc/DOSXYZnrc for a 10 × 10 cm2 field versus measurement

with a diamond at depths of 1.4, 10 and 20 cm. b) Absolute difference

between calculation and measurement for each depth. .............................. 104

Figure 6-3 Percentage depth doses calculated by BEAMnrc/DOSXYZnrc and

measured with a diode for: a) 4 mm circular collimator, b) 10 mm

circular collimator, c) 20 mm circular collimator and d) 30 mm

circular collimator. ..................................................................................... 105

Figure 7-1 In-house Virtual Water phantom showing a low density slab

inserted in the middle section .................................................................... 119

Figure 7-2 Workflow diagram showing the process followed for comparing

iPlan generated treatment plans with Monte Carlo generated plans

using MMCTP platform. ............................................................................ 121

Figure 7-3 Anthropomorphic phantom showing treatment plan with 7.5mm

Brainlab cone ............................................................................................. 123

Figure 7-4 Percentage depth dose calculated by iPlan and by Monte Carlo for

the Brainlab circular cones in a virtual water phantom containing a

low density region. ..................................................................................... 124

Figure 7-5 Percentage depth dose calculated by iPlan and by Monte Carlo for

the Brainlab circular cones in a virtual water phantom containing a

high density region. .................................................................................... 125

Figure 7-6 Cross profiles at depth calculated by iPlan and by Monte Carlo for

the Brainlab circular cones in a slab of low density material inserted in

virtual water phantom ................................................................................ 126

Figure 7-7 Cross profiles at depth calculated by iPlan and by Monte Carlo for

the Brainlab circular cones in a slab of high density material inserted

in virtual water phantom ............................................................................ 126

x Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

Figure 7-8 Dose Volume Histogram for the dose distribution calculated in an

anthropomorphic phantom using iPlan and Monte Carlo for a 7.5 mm

and 10 mm cone. ........................................................................................ 128

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery xi

List of Tables

Table 2-1 Penumbra (80%-20%) and FWHM measurements by an IBA SFD

diode and a PTW 60019 microDiamond detector........................................ 33

Table 2-2 Field factors obtained using a PTW 60012 E diode, an IBA SFD

diode and PTW 60019 microDiamond detector at a depth of 1.4 cm

for a 6 MV SRS x-ray beam on a Novalis Tx equipped with circular

cones at an SSD of 100 cm. The uncertainties were up to 0.5% (1 SD)

for all detectors. ........................................................................................... 34

Table 2-3 Monte Carlo calculated correction factors for PTW 60019

microDiamond detector at a depth of 1.4 cm for a Novalis Tx

equipped with circular cones using a 6 MV SRS x-ray beam ..................... 35

Table 3-1 Monte Carlo calculated correction factors,

kQclin, Qmsrfclin, fmsr[sfd], for Gafchromic EBT3 film for a 6 MV

SRS beam ..................................................................................................... 50

Table 4-1 Measured and extrapolated relative output factors for Gafchromic

EBT3 film with different analysis diameters and for the PTW 60019

microDiamond detector ............................................................................... 67

Table 4-2 Analysis of detector effective size versus pixels measured for various

resolutions .................................................................................................... 67

Table 4-3 The DPI scanning resolution used across studies for small field

dosimetry...................................................................................................... 69

Table 6-1 Relative surface doses for Brainlab SRS circular collimators

determined by Monte Carlo calculations and Gafchromic EBT3

measurements ............................................................................................. 106

Table 7-1 FWHM90-10 and FWHM80-20 values for the High and Low density

slab materials for the 4, 7.5 and 10 mm cones. .......................................... 127

xii Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

List of Publications for PhD Candidature

1. J.E. Morales, S.B. Crowe, R. Hill, N. Freeman, J.V. Trapp “Dosimetry of

cone-defined stereotactic radiosurgery fields with a commercial synthetic

diamond detector” Medical Physics, 41, 2014.

2. J.E. Morales, M. Butson, R. Hill, S.B. Crowe, J.V. Trapp “Monte Carlo

calculated output correction factors for Gafchromic EBT3 film for dosimetry

in small stereotactic radiosuregry fields” Submitted for publication 2019 in

APSEM Journal.

3. J.E. Morales, M. Butson, S.B. Crowe, R. Hill, J.V. Trapp “An experimental

extrapolation technique using Gafchromic EBT3 film for relative output factor

measurements in small x-ray fields” Medical Physics, 43, 2016.

4. T. P. Huang, J.E. Morales, E. Butson, A. Johnson, M. Butson, Robin Hill, “A

novel extrapolation method using OSL detectors for very small field output

factor measurement for stereotactic radiosurgery” Submitted for publication

2019 in APSEM Journal.

5. J.E. Morales, R Hill, S. B. Crowe, T. Kairn, J. V. Trapp, “A comparison of

surface doses for very small field x-ray beams: Monte Carlo calculations and

radiochromic film measurements” Australas Phys Eng Sci Med (2014) 37:303–

309

6. J.E. Morales, M. Butson, R. Hill, S. B. Crowe, J.V. Trapp “A study of dose

inhomogeneity correction in a commercial treatment planning system for

stereotactic radiosurgery” Submitted for publication 2019 in APSEM Journal.

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery xiii

List of Other Research Activities and Published

Works during PhD Candidature

1. I. Das. J. E. Morales, P. Francescon “Small field dosimetry: What have we

learnt?” AIP Conference Proceedings 1747, 060001 (2016); doi:

10.1063/1.4954111.

2. I.Das, Textbook: “Radiochromic film. Roles and Applications in Radiation

Dosimetry”. (2017) PhD candidate J.E. Morales co-authored Chapter 13

entitled: “Small field dosimetry in megavoltage beams” This chapter was based

on the work in Chapter 4 of this PhD thesis. Textbook doi:

https://doi.org/10.1201/9781315154879

3. Morales, J.E., Hill, R., Crowe, S., & Trapp, J. (2014) “Commissioning of a

new commercial plastic scintillator system for radiotherapy”. Australasian

Physical and Engineering Sciences in Medicine, 37(1), p. 177.

4. Butson M, Haque M, Smith L, Butson E, Odgers D, Pope D, Gorjiana T,

Whitaker M, Morales J, Hong A, Hill R. (2017) “Practical time considerations

for optically stimulated luminescent dosimetry (OSLD) in total body

irradiation”. Australas Phys Eng Sci Med. 2017 Mar;40(1):167-171.

doi:10.1007/s13246-016-0504-4.

5. Smith L, Haque M, Morales J, Hill R. “Radiation dose measurements of an

on-board imager X-ray unit using optically-stimulated luminescence

dosimeters”. Australas Phys Eng Sci Med. 2015 Dec;38(4):665-9.

6. Morales, Johnny, Crowe, Scott, & Trapp, Jamie (2012) “Monte Carlo

modeling of a 4 mm conical collimator for a Novalis Tx Linear Accelerator”.

In EWGMCTP - Third European Workshop on Monte Carlo Treatment

Planning, 15-18 May 2012, Seville, Spain.

https://doi.org/10.1201/9781315154879

xiv Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

Statement of Original Authorship

The work contained in this thesis has not been previously submitted to meet

requirements for an award at this or any other higher education institution. To the best

of my knowledge and belief, the thesis contains no material previously published or

written by another person except where due reference is made.

QUT Verified Signature

Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery xv

Acknowledgements

I would like to thank and acknowledge my supervisors, Dr Jamie Trapp and Dr

Scott Crowe, for allowing me to undertake this degree under their supervision. I thank

them for sticking with me till the end despite the obstacles that I faced in the beginning.

It is important to note that I undertook this project while I was working full time so

much patience was needed in helping me to progress with the publications. Their

support at the university is invaluable and while not readily evident it went a long way

in helping me to make progress one step at a time. Their loyalty to me is something

that I will never forget.

I would like to also thank my colleagues Dr Robin Hill and Dr Martin Butson.

Robin is my manager and made decisions that helped me develop my professional

career especially in the field of SRS. Robin assigned me to take care of the SRS

program in our Department and that enabled me to be in touch with developments in

the field. Robin also facilitated most, if not all, of my conference attendances in

Australia and Overseas. Martin came in later in the project and was tremendous source

of advice in radiochromic film research and the publications of a number of papers.

I am also very thankful to two of my colleagues who helped me look after the

SRS program in my Department at the beginning of my candidature. They were Simran

Gill (MSc) and Dr Samara Alzaidi, their tireless work, great support and enthusiasm

in maintaining the clinical SRS work at the beginning of my project was invaluable.

I also would like to thank Dr Tanya Kairn and Nigel Freeman (MSc) who were

part of my papers published during this project.

I would to thank the Director of Radiation Oncology at Chris O’Brien Lifehouse,

Associate Professor Dr Chris Milross, for his constant support for our Physics team

CPD activities. Dr Milross has always ensured that we get support and this has made

a difference to our team development. It has made possible for me to attend

conferences and workshops in Australian and overseas, thus ensuring my network

development with key researchers around the world.

xvi Advances in very small x-ray field dosimetry for circular cones used in stereotactic radiosurgery

Finally, I thank my parents and family for their unconditional support throughout

these years. I will always appreciate their tireless support and encouragement. Their

belief in me has always been a source of strength.

Chapter 1: Introduction 1

Chapter 1: Introduction

1.1 RADIATION THERAPY AND STEREOTACTIC RADIOSURGERY

1.1.1 CANCER IN AUSTRALIA

Cancerous tumours are those where malignant cells attack healthy cells in a

localised body region1, which can be lethal if vital organs are attacked and if malignant

cells proliferate to other parts of the human body2. The Australian Department of

Health and Aging estimates that cancer is still the one of the major diseases in the

Australian population3 and that in 2018 there will be more than 130000 newly

diagnosed cancers in Australia. In 2018 cancer accounted for about 19% of the total

disease burden in the public health system and it is estimated that on average about

48000 people die from cancer every year3. It is estimated that the Australian

government provided over $10 billion dollars in funding for cancer control activities

alone in the period from 2014 to 2017.

Cancer treatments can make use of a range of modalities including surgery,

chemotherapy, radiation therapy and immunotherapy4-6. For many types and stages of

cancer, the patient may receive a combination of these treatments7. For some cancers,

radiation may be used to shrink the tumour before the patient has surgery or may be

used after the surgery to reduce the risk of the cancer reoccurring8, 9. In some other

forms of cancer, a few cycles of chemotherapy may be given before the radiation

therapy starts10.

Radiation therapy makes use of ionising radiation to destroy the cancer cells by

causing breaks in the DNA of the cells 1. The ionising radiation used in radiation

therapy can be X-rays, electrons, protons or higher atomic number ions; with X-rays

being most common11.

One of the key aims of radiation therapy is to maximise the radiation dose to the

tumour while minimising the dose to the surrounding tissues and organs at risk nearby

the tumour12. This is achieved by projecting the radiation beams into the tumour and

the process of fractionation where the total radiation dose is delivered in many

fractions of smaller doses over a number of weeks12. This fractionation ensures that

2 Chapter 1: Introduction

maximum cell kill in the tumour occurs while also allowing the healthy tissues to

recover and repair from the radiation13.

It is estimated that approximately 40% of cancer cures are attributed to radiation

therapy14 and there is evidence that nearly 48% of patients with cancer in Australia

would benefit from radiation therapy at some stage during the course of their illness15.

Some of the main types of cancer which are treated with radiation therapy either as a

cure, as palliative treatment to provide symptom relief or as adjuvant therapy are :

lung cancer16, breast cancer17, prostate cancer18, 19, cervical cancer20, colon cancer21,

kidney cancer22, liver cancer23, pancreatic cancer24, rectal cancer25, stomach cancer26,

oesophageal cancer7 and brain cancer7. Therefore, the role that radiation therapy plays

in the fight against cancer is of paramount importance.

1.1.2 RADIATION THERAPY DELIVERY

Modern radiation therapy using external x-ray beams is typically delivered by

an electron linear accelerator (linac)11 . These radiotherapy linacs accelerate electrons

to high kinetic energies, travelling at almost relativistic speed11 where they strike a

target and produce bremsstrahlung11. This bremsstrahlung are high energy x-rays

which are then directed to the tumour in the patient. There are several manufacturers

of linacs, with the most widely available being Varian (Varian Medical Systems, Palo

Alto, USA) and Elekta (Elekta Instruments AB, Stockholm, Sweden). These devices

are capable of producing x-rays with energies of up to 18 Megavolts. Figure 1 shows

a Novalis Trilogy linear accelerator (Varian Medical Systems, Palo Alto, USA).


Figure 1-1 Novalis Tx linear accelerator capable of producing 6 MV x-rays for

treatment in radiotherapy

1.1.3 WHAT IS STEREOTACTIC RADIOSURGERY

The term stereotactic refers the three-dimensional localisation of a particular

point in space relative to a fixed external frame27. The frame acts as a support for

hollow probes carrying electrodes or biopsy needles to precise locations within the

human brain based on orthogonal x-ray film of the brain in the frame27. This principle

has been used since 195128 by neurosurgeons28 for stereotactic radiosurgery (SRS)

which utilises narrow beams of 60Co gamma photons focussed on a small target within

the human brain. The initial aim in the development of SRS was to provide a “non-

invasive destruction of intracranial lesions which may be unsuitable for open

surgery”28. Initially, treatments used orthovoltage X-ray units to treat trigeminal

neuralgia, however, Leksell ultimately designed a Gamma Knife unit which became

operational in 1968. The Gamma Knife includes four collimators measuring 4, 8, 14

and 18 mm in diameter.

Over time, other groups around the world started developing other options for

delivery of SRS treatments using linacs29, 30. There have been recent advancements

which allow linacs to provide stereotactic radiosurgery treatments to lesions in the

brain. One of those technological advancements has been the invention of multi-leaf

collimators31 (MLCs) which modify the cross-section shape of the treatment beam to

match the shape of the lesion. MLCs come in different sizes, with some models as

narrow as 2.5 mm in width while others can be up to 5 mm wide. The choice of size


depends on the clinical need. For SRS treatments, smaller widths are generally

required32.

In addition to MLCs, circular collimators or circular cones are also used for SRS

treatments1. Circular cones, as the name implies, produce circular shaped x-ray fields.

They are usually tertiary devices in that they are attached to the linac head as an add-

on device and are typically in sizes ranging from 4 mm up to 30 mm in diameter. For

SRS treatments, the most widely used circular cones range from 4 to 10 mm although

bigger sizes are sometimes used. SRS treatments using circular cones on a linac are

usually delivered as isocentric arcs, rotating through a range of degrees where the

isocentre is usually at the centre of the lesion. Figure 1.2 below shows a set of Brainlab

circular cones used for SRS treatments on a linear accelerator.

Figure 1-2 A set of Brainlab circular cones used for SRS treatments on a linear

accelerator.

While the treatments with a Gamma Knife required the use of a frame, the SRS

treatments using MLCs or circular cones on a linac can be delivered on a frame-less

environment33, 34, offering a much improved patient experience. Advancements in on-


board imaging technology on Linacs has made possible the improvement of patient set

up via digital x-ray imagining using technologies such as On-Board Imaging (Varian,

Palo Alto, USA) and ExacTrac Brainlab Technology (Brainlab, Germany)34, 35. In a

frameless environment usually a thermoplastic mask is used and customised to fit the

patient specific cranial contour36.

In addition to being used for treating very small brain cancers, the smaller

diameter Brainlab cones can also be used as one of the treatment options for trigeminal

neuralgia 37. Trigeminal neuralgia is a condition that that affects the trigeminal nerve

and results in very chronic pain for the patient 38-41. The trigeminal nerve carries

sensations between the brain and the face. The treatment for trigeminal neuralgia

involves a stereotactic radiosurgery technique using very small fields, typically using

field sizes of 4 or 5 mm diameter 39, 42. Prescribed doses reported in the literature for

these treatments vary from 60 Gy up to more than 90 Gy which is delivered in a single

treatment fraction 43-45. Therefore with these very high radiation doses being delivered,

accurate small field dosimetry is also required for these treatments 42.

1.2 DOSIMETRY IN RADIOTHERAPY

1.2.1 Reference Dosimetry – Absorbed Dose

The absorbed dose is defined as the mean energy (Joules) imparted to a point of

mass (Kg). The unit of absorbed dose is the Gray where 1 Gray is equivalent to 1 J/Kg2.

To ensure consistency in reference dosimetry around the world, there are international

codes of practice (COP). These COP provide guidelines on the measurement

techniques, choice of detectors, correction quantities and more importantly a reference

framework which can be followed by clinical department At the moment, the two main

COPs that are in use are the IAEA TRS-39846 and AAPM TG – 5147. Both COPs have

been designed by metrologists and other experts in the field of dosimetry and form the

basis of international best practice for most hospitals. These codes provide guidelines

for reference dosimetry of high and low energy photon beams, electron beams and

particle beams such as protons.

1.2.2 Dosimeters used for Reference Dosimetry

The reference detector currently recommended for determining absolute

absorbed dose to water for photon beams is the thimble-type ionisation chamber11.

These detectors have a cylindrical shape typically with volumes from 0.1 – 1.0 cm3


and have a thin wall of graphite or plastic. The outer shell is usually operated at ground

potential and the central electrode conducts the applied voltage. The ionisation

chambers are designed to collect ion pairs created in the small air cavity by the

interaction of the ionising radiation beams. One of the main qualities of the ionisation

chamber is the its robustness and ideal shape for taking measurements under electronic

equilibrium conditions in photon beams which makes these detectors achieve great

reproducibility and accuracy during such measurements.

According to the IAEA TRS-398 COP46, the recommended geometric set up for

determining the reference dose to water using an ionisation chamber are: a field size

of 10 × 10 cm2, Source to Axis distance of 100 and depth of 10 cm. However, there is

allowance for performing this measurement at a source to Surface Distance of 100 cm

keeping the field size and depth the same. The Monitor Units (MUs) used for this

measurement are usually 100 MUs and the output of the linac is then quoted as 1

cGy/MU.

1.2.3 Relative Dosimetry

Relative dosimetry involves the measurement of dose for the photon beam in

any conditions other than the reference conditions as specified in the COP. The word

relative is used because the value of dose measured is compared to the reference

conditions for the radiation device being used. If the radiation device is a linear

accelerator then the usual reference conditions in Australia can be a field size of 10 ×

10 cm2, source to surface distance of 100 cm and a depth of 10 cm.

There are various detectors which are currently used for relative dosimetry in

photon beams. Some detectors can be used for large beams but other detectors are

specially designed for measurements of small fields. For the purpose of this thesis, the

term small refers to fields less than or equal to 30 mm diameter. The paragraphs below

present a brief description of some of the detectors used for the work in this thesis,

these detectors are silicon diodes, microDiamond, radiochromic film and optically

stimulated luminescence dosimeters (OSLDs).

1.2.4 Relative Dosimetry - Diodes

Silicon diode detectors are one of the most widely used semiconductors for

radiation dosimetry. The sensitive region in a silicon diode is the p-n junction48. These

detectors don’t need the application of any voltage to create an electric field. During


irradiation by the beam, electron-hole pairs are created and the charge carriers are

swept across the junction by an intrinsic potential.

Diode detectors can be used in either current or charge mode. For this type of

detector, when used in current mode, the current induced by the radiation is

proportional to the dose rate. When used in integral mode the total charge is

proportional to the total dose measured. Silicon diodes can be manufactured to very

small sizes which makes them ideal for measurements in small x-ray fields. However,

diodes have some limitations that include having an overresponse to low energy

photons and an overresponse11 for fields greater than 10 × 10 cm2. In addition, their

response for very small fields can also vary significantly and so corrections may need

to be applied for measurements in fields less than or equal to 30 mm diameter49.

1.2.5 Relative Dosimetry – diamonds and synthetic diamonds

Natural diamond detectors have been used in radiotherapy for many years50.

They are semiconductors and can be either p- or n-type devices48. Diamond detectors

have a high sensitivity and have a constant ratio of water to diamond stopping power

ratio as well as mass-absorption coefficients. It is these particularly qualities that make

them very good for performing dose measurements in water. However, they also have

a number of disadvantaged including having variable response according to the dose

rate of the radiation beam. In addition, Diamond detectors need to be given a

substantial pre-irradiation dose in order for their signal to stabilise48. They require no

bias potential similar to silicon diodes when taking measurements. However, with all

of these advantages and disadvantages being considered, the main issue for diamond

detectors was the very high cost and for this reason they are no longer being

manufactured.

In recent years, natural diamonds have been replaced by artificial diamonds

which are manufactured through a chemical vapour deposition technique (CVD)51.

These artificial diamonds can be manufactured to very small dimensions and to a very

reproducible degree of accuracy. This has generated a lot of interest in the radiotherapy

community with possible applications with a wide range of beams. They do have their

particular disadvantages and their high density (3.52 gcm-3) relative to water can be

one of them.


1.2.6 Relative Dosimetry – Radiochromic film

Radiochromic film works on the principle of colouration caused by ionising

radiation48. The degree of coloration is proportional to the amount of dose deposited.

The colouration is the result of polymerisation of the special dye contained in the

sensitive part of the film48. Radiochromic film have a number of advantages starting

with that this type of film is self-developing and does not need any special developing

including the use of chemicals. These films are not sensitive to ambient light although

there is some sensitivity to Ultra Violet light. Radiochromic film is regarded as being

near tissue equivalent for the megavoltage range of photon energies. It has high spatial

resolution and can be very useful to measure 2D dose maps and can be used for

measurements in small x-ray fields. However, the process to read out radiochromic

film is challenging and requires extreme care to achieve good reproducibility52.

1.2.7 Relative dosimetry – Optically Stimulated Luminescence Dosimeters

These detectors work on the principle of optically stimulates luminescence and

hence their name of optically stimulated luminescence dosimeters. These detectors are

made up of crystalline dielectric material (Al2O3), aluminium oxide, with added

contaminants (C) which form a imperfections in the crystal structure53. These

contaminants play the role of traps for electrons or holes produced by radiation. When

the detector is optically stimulated then electrons are ejected out of these traps and

light is emitted. The intensity of the luminescence is proportional to the dose measured

by the crystal. The main advantage of OSLDs is their ease of use and readout process.

Their disadvantage is their size for small field dosimetry is their size. They are also

non-tissue equivalent due to their composition being aluminium oxide. The

commercially available OSLDs are 7 mm in diameter which could potentially be too

large for some of the very small fields used in this thesis.

1.3 LITERATURE REVIEW – WHAT HAVE WE LEARNT?

1.3.1 ADVANCES IN RADIATION TECHNOLOGY

Evolution in technology has changed radiation therapy to a high degree of

sophistication and complexity. Advances in stereotactic radiosurgery54, 55 (SRS),

stereotactic body radiotherapy56 (SBRT), for cranial and extra-cranial lesions and

intensity modulated radiation therapy (IMRT) and volumetric modulated radiotherapy

(VMAT) use relatively small fields (<3 cm) that can be either static or dynamic57, 58.


This has created many innovations in treatment machines such as various designs of

Gamma Knife and linear accelerators that deliver relatively small fields either in

specialized cones, iris, or multileaf collimators (MLC). Traditional radiation oncology

fields that are commissioned span from 3×3 cm2 to a maximum of typically 40×40 cm2

as described by Das et al 59. Dosimetry in small fields is a complicated and relatively

new field 60.

1.3.2 SMALL FIELD DOSIMETRY CHALLENGES

Using these small beams does come with its own challenges. Using broad

radiotherapy beams (as traditionally used) the dose is well understood and predictable.

As the size or shape of the radiation field is varied, the radiation dose at a particular

depth also varies due to contributions from secondary radiation which has been

scattered in the medium. Simply put, a larger field will have a larger scatter

contribution and will result in a higher relative dose. Tabulated data exists for

calculating the dose in relation to a reference field size, which is typically 10 × 10 cm2.

Once a radiotherapy field is reduced to smaller than about 1.5 cm in size, the

previous theories, assumptions and methodologies for calculating dose begin to break

down 61-63. Some of the known contributing factors include a loss of lateral charged

particle equilibrium at 1.5 cm 61, 62, 64 which is primarily responsible for the ‘dose to

field-size dependence’ of broad fields as discussed above, penumbral overlaps 64 and

penumbral dominance, spectral changes due to the beam transiting only the central

portion of the accelerator’s flattening filter 64, and source occlusion below 8 mm 61. In

addition, advanced and developing radiotherapy modalities such as intensity

modulated radiotherapy (IMRT), volumetric modulated arc therapy (VMAT),

stereotactic body radiotherapy (SBRT), depend on modulated fields and require

similar small field dosimetry methods to fully characterise them.

Further adding to the uncertainties associated with small fields are a number of

factors that research has shown to be specific to each individual linear accelerator.

Recent research has shown that:

• The radiation field that is delivered to the patient is different for the same

operator settings across individual linear accelerators of the same model 65, 66.

• For each linear accelerator there are specific contributions to small field

uncertainties from a variety of sources such as x-ray scatter from the


collimating jaws and the housing of the head of the accelerator, occlusion of

the focal spot of the electron beam, and photon scatter within the patient 61.

Furthermore, there are difficulties in accurate measurement of small fields. For

a typical radiation detector to work correctly the entire volume must be encompassed

within the field to avoid volume averaging effects and to meet design criteria of the

detector. E.g. , the ‘workhorse’ detector for radiotherapy has traditionally been

ionization chambers using the Bragg-Gray cavity theory (BGCT) 67 and its subsequent

revisions 68-70. However, the accuracy of ionization chambers with a small enough

volume for small fields have come into question 71 leading to investigation of the

suitability and potential use of different detector types 71-75.

The difficulty of dosimetry in small fields lies in the electron transport created

by the photon interaction with medium. Das et al76 and Charles et al77 have provided

definition of small fields that depends on dose disequilibrium, source size and more so

the selection of detector. In the past SRS dosimetry had been uncertain for small fields

by as much as 14% among institutions and detectors 78. Typically published data from

major institutions have been used as gold standard for dedicated devices 79 however

this had large errors. The dosimetry protocols like the IAEA TRS-398 80 and AAPM

TG-51 81 provided guidelines for reference field size which is typically 10 × 10 cm2.

Most reference conditions parameters such as stopping power ratio, perturbation

correction, fluence and gradient corrections are not applicable to small fields. To

overcome non-reference fields by specialized machines, the International Atomic

Energy Agency (IAEA) has provided a framework of an international approach to deal

with the issues in small field dosimetry 82. In the same time frame, AAPM formed a

task group (TG-155) to provide relative dosimetry in small fields 83.

1.3.3 THE INTERNATIONAL ATOMIC ENERGY AGENCY (IAEA)

APPROACH TO SMALL FIELD DOSIMETRY

In 2017, the IAEA and the American Association of Physicists in Medicine

(AAPM), jointly released a new and more robust Code of Practice, TRS-48349, entitled

Dosimetry of Small Static Fields Used in External Beam Radiotherapy – An

International Code of Practice for Reference and Relative Dose Determination. This

new international Code of Practice (COP) provides guidance and recommendations

for both relative and reference dosimetry using small x-ray fields. It provides specific


advice on how to perform reference dosimetry in non-standard machine specific

reference (fmsr) fields.

In the new TRS-483 formalism, the mathematical expression for absorbed dose

to water, 𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 , for a machine specific reference field, fmsr, is given by :

𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 = 𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 𝑁𝐷,𝑤,𝑄𝑜

𝑓𝑟𝑒𝑓 𝑘𝑄𝑚𝑠𝑟,𝑄𝑜

𝑓𝑚𝑠𝑟,𝑓𝑟𝑒𝑓

Where 𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 is the reading of the dosimeter in the machine specific reference

field, fmsr, corrected for influence quantities, 𝑁𝐷,𝑤,𝑄𝑜

𝑓𝑟𝑒𝑓 is the calibration coefficient in

terms of absorbed dose to water measured at a standards laboratory for a conventional

10 × 10 cm reference calibration field, fref , with beam quality Qo, 𝑘𝑄𝑚𝑠𝑟,𝑄𝑜

𝑓𝑚𝑠𝑟,𝑓𝑟𝑒𝑓 is the

factor that corrects for the difference in the response of the ionization chamber in a

conventional 10 × 10 cm reference field, fref, with beam quality Qo using the same

machine as the machine specific reference field, fsmr, and the response of the ionization

chamber in the machine specific reference field fmsr with beam quality Qmsr.

Furthermore, the new IAEA TRS-483 COP provided guidance for measurements

of field output factors and lateral beam profiles at the measurement depth because of

their importance in the determination of the field size and the volume averaging

correction of the particular detector used. When a field size falls below the lateral range

of charged electronic equilibrium, the output factor measurements obtained with

certain radiation detectors may need correction applied to its measurement.

The new COP defines the output factor, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , relative to machine specific

reference field, fmsr, and for small fields as:

Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 =𝑀𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛

𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟

Where 𝑀𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛 and 𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 are the readings of the detectors in the clinical field,

fclin, and machine specific reference field, fmsr, respectively. The values for the output


correction factor 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 are provided in Tables 23 to 27, found in the COP, for

different x-ray beam energies and different technology configurations such as linear

accelerators, the CyberKnife unit and Tomotherapy unit all for a wide range of

commercially available detectors.

These 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 factors were derived from a subset of published data available

at the time of the report using experimental methods and Monte Carlo calculations.

This data was determined to be necessary in order to provide a consistent methodology

to determine field output factors for small fields. The adoption of the TRS-483 COP

is now assisting the Medical Physics community worldwide in standardising the

methodology for small field dosimetry84, 85.

1.4 AIMS OF THE PROJECT

At the start of this project in 2012, there were many gaps in the scientific

knowledge of small field x-ray dosimetry, which impacted both the clinical and

research environment. For example, what are suitable detectors that can be used in the

clinic for accurate relative dosimetry measurements of Brainlab cones? While the

vendor provided a methodology for the measurements, they did not provide either

recommended detectors or reference beam data for comparison.

Therefore, one of the aims of this project was to determine detector suitability

for clinical use and application in small field x-ray dosimetry for very small fields, as

small as 4 mm in diameter. Several detectors were compared, including the newly

released PTW microDiamond 60019 detector from PTW-Freiburg. At time of this

work, there was no literature available on its characteristics, particularly for small

fields as defined by stereotactic radiosurgery cones86, 87. It should be noted that this

work was performed and published before the release of the new IAEA Code of

Practice, TRS-48349. This new COP now provides data for the microDiamond detector.

There were further gaps in the knowledge of detector suitability particularly for

2D dosimetry measurements of small x-ray fields. A number of earlier publications88,

89 had investigated the previous models of radiochromic film (Gafchromic EBT2

version) as a reference dosimeter but this had not extended to the currently available

EBT3 film90. While radiochromic film has a number of positive characteristics for

small field dosimetry, including high spatial resolution and radiological equivalence


to water, its inherent performance for small field dosimetry had not been investigated,

and thus characterisation was undertaken in this project with the Gafchromic EBT3

film91.

In addition, there were gaps in the knowledge of how can film dosimetry as a

technique be optimised for the most accurate dosimetry of small fields, leading to two

further studies in this work91. The concept to optimise the methodology to remove the

volume averaging effect of a detector was extended to OSLDs detectors by reducing

the area of the detector. This novel method not only showed how to extend the

extrapolation technique to other detectors but it also conveyed the idea of

manufacturing new OSLDs detectors for small field dosimetry.

There were also gaps in the knowledge in clinical aspects of SRS treatments with

Brainlab cones. Firstly, the surface (skin) doses were unknown, which was

investigated and published92. Secondly, the accuracy of the treatment planning system

was up to 30% for single beam geometry and thus an investigation was undertaken

which showed that the combined inaccuracy can be up to 12% for multiple beam

arrangements.

1.5 THESIS OUTLINE

In this thesis, the work has been compiled in a number of publications that fill in

gaps in knowledge and each contribute to the overall understanding and advancement

in small field dosimetry from Monte Carlo modelling of a Novalis Tx linear accelerator

equipped with Brainlab circular cones to dosimetry of small field beams produced by

these circular cones.

Chapter 2 - Dosimetry of cone-defined stereotactic radiosurgery fields with a

commercial synthetic diamond detector. This paper was one of the first studies

published on the performance of a new commercial synthetic diamond detector being

the PTW 60019 microDiamond from PTW-Freiburg. Relative dosimetry

measurements with the microDiamond were performed for small x-ray fields down to

4 mm diameter. These dose values were compared with both Monte Carlo calculations

as well as dosimetry data measured using diodes detectors that were considered gold-

standard at that time.

Chapter 3 - Monte Carlo calculated output correction factors for Gafchromic

EBT3 film for dosimetry in small stereotactic x-ray fields. This paper presented the


first study which has calculated small field output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for

Gafchromic EBT3 film. Monte Carlo calculations were used to model the film and

determine whether these correction factors can be assumed to be unity or not as has

been alluded to in literature to date.

Chapter 4 - An experimental extrapolation technique using Gafchromic EBT3

film for relative output factor measurements in small x-ray fields. This paper outlines

a novel extrapolation technique to determine relative output factors for very small x-

ray fields using Gafchromic EBT3 film. This technique involved plotting the relative

output factors as a function of the ROI used in the analysis. From this, the relative

output factor was extrapolated to a zero area to determine the final relative output

factor.

Chapter 5 - A novel extrapolation method using OSL detectors for very small

field output factor measurement for stereotactic radiosurgery. The work in this paper

involves adapting the extrapolation technique from Chapter 4 to the NanoDot which

is an optically stimulated luminescent dosimeter (OSLD). Commercial OSLD

detectors were physically modified to give a different effective area for the readout.

The final relative output factor was determined by a similar extrapolation technique to

a zero-effective area.

Chapter 6 - A comparison of surface doses for very small field x-ray beams:

Monte Carlo calculations and radiochromic film measurements. To date, there is

limited data on surface or skin doses for very small x-ray fields. The study presented

in this paper involved experimental measurement of surface doses using radiochromic

film as well as a full Monte Carlo model of the respective x-ray beams. This was the

first known publication that examined surface doses for field sizes less than 5 mm.

Chapter 7 - A study of dose inhomogeneity correction in a commercial treatment

planning system for stereotactic radiosurgery. This paper presents a pilot study using

Monte Carlo methods to investigate the accuracy of a stereotactic radiosurgery

treatment planning system in modelling tissue inhomogeneities. The focus of the study

was for very small field sizes of 4.0, 7.5 and 10.0 mm diameter. Dose calculations in

the treatment planning system were calculated in block phantoms containing tissue

inhomogeneities as well as in an anthropomorphic head phantom.


Finally, Chapter 8 gives a summary of the findings of the research in this thesis

as well as giving suggestions for future work in small field dosimetry particularly as

applied to stereotactic radiosurgery.


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23

Statement of Co-Authors for Chapter 2



Chapter 2: Dosimetry of cone-defined stereotactic radiosurgery fields with a commercial synthetic diamond

detector 25

Chapter 2: Dosimetry of cone-defined

stereotactic radiosurgery fields

with a commercial synthetic

diamond detector

Overview

Natural diamond detectors have been considered excellent detectors for radiation

dosimetry. However, they are hard to manufacture to reproducible volume dimensions.

In the past decade, there has been significant improvement in the development of

artificial diamond detectors for radiation dosimetry. These artificial diamonds are

grown by a process of chemical vapour deposition (CVD). A new commercial

synthetic diamond detector came into the market in 2014. This detector was based on

CVD technology. The official name of the detector was PTW 60019 microDiamond.

The objective of this paper was to evaluate the performance of this new commercial

synthetic diamond detector for the dosimetry of small x-ray fields as used for

stereotactic radiosurgery on a Novalis Tx linear accelerator equipped with circular

cones. Correction factors were provided in this paper, in 2014, for a geometric set up

corresponding to the method recommended by Brainlab for use in the treatment

planning system used for planning with the small circular cones. In 2014, there was no

official recommended methodology to determine the output factors for small fields

used in stereotactic radiosurgery (SRS). However, a new recommended method to

determine correction factors was introduced in 2017 by the IAEA as published in the

Code of Practice TRS 483. The main difference was in the depth of measurement. The

IAEA recommended a depth of 10 cm while this paper used 1.5 cm as recommended

for Brainlab. The main results show that the IAEA had correction factor of 1.04 while

the paper in this thesis had a correction of 1.027. Furthermore, the paper in this chapter

provided recommendations for cross profiles measurements in water.

26 Chapter 2: Dosimetry of cone-defined stereotactic radiosurgery fields with a commercial synthetic diamond

detector

STATEMENT OF JOINT AUTHORSHIP

Title: Dosimetry of cone-defined stereotactic radiosurgery fields with a

commercial synthetic diamond detector

Authors: Johnny E Morales, Scott B Crowe, Robin Hill, Nigel Freeman, J.V.

Trapp

Johnny E Morales (candidate)

Performed all measurements and Monte Carlo calculations. Involved in the

project design and wrote the entire manuscript.

Scott B Crowe

Provided advice and supervision as required. Helped with interpretation of

results. Provided feedback on manuscript write up.

Robin Hill

Helped with interpretation of results. Provided feedback on manuscript write up.

Nigel Freeman

Helped with interpretation of results and provided feedback on the project.

J.V. Trapp

Supervised the project and provided direction. Helped with interpretation of

results. Edited manuscript and contributed to the write up.

Journal: Medical Physics

Status: Published 2014

SCOPUS Citations to date: 40

SCOPUS Authors h-index: 6

http://dx.doi.org/10.1118/1.4895827


detector 27

ABSTRACT

Purpose: Small field x-ray beam dosimetry is difficult due to a lack of lateral

electronic equilibrium, source occlusion, high dose gradients and detector volume

averaging. Currently there is no single definitive detector recommended for small field

dosimetry. The objective of this work was to evaluate the performance of a new

commercial synthetic diamond detector, namely, the PTW 60019 microDiamond, for

the dosimetry of small x-ray fields as used in stereotactic radiosurgery (SRS).

Methods: Small field sizes were defined by Brainlab circular cones (4 – 30 mm

diameter) on a Novalis Trilogy linear accelerator and using the 6 MV SRS x-ray beam

mode for all measurements. Percentage depth doses (PDDs) were measured and

compared to an IBA SFD and a PTW 60012 E diode. Cross profiles were measured

and compared to an IBA SFD diode. Field factors, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were calculated by

Monte Carlo methods using BEAMnrc and correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were derived

for the PTW 60019 microDiamond detector.

Results: For the small fields of 4-30 mm diameter, there were dose differences in the

PDDs of up to 1.5% when compared to an IBA SFD and PTW 60012 E diode detector.

For the cross profile measurements, the penumbra values varied, depending upon the

orientation of the detector. The field factors, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were calculated for these field

diameters at a depth of 1.4 cm in water and they were within 2.7% of published values

for a similar linear accelerator. The corrections factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were derived for

the PTW 60019 microDiamond detector.

Conclusions: We conclude that the new PTW 60019 microDiamond detector is

generally suitable for relative dosimetry in small 6 MV SRS beams for a Novalis

Trilogy linear equipped with circular cones.


detector

2.1 INTRODUCTION

Stereotactic radiosurgery (SRS) involves the delivery of a high radiation dose to

lesions within the brain using small field size radiation beams.1-3 The dosimetry of

very small x-ray fields is challenging for several reasons including a lack of lateral

electronic equilibrium, source occlusion, large dose gradients and the size of detector

in respect to the field size.4-6 There have been many investigations into the choice of

appropriate radiation dosimeters for relative dosimetry measurements such as depth

doses, profiles and relative output factors in very small x-ray fields.7-9 The detectors

studied have included very small ionisation chambers (pinpoint chambers), diodes,

diamond detectors, plastic scintillator dosimeters (PSDs) and radiochromic film.7, 10, 11

The incorrect choice of detector can result in up to 30% difference in relative output

factor leading to radiation accidents and the need for significant correction factors have

been reported particularly for very small field sizes12-14.

Recently, there has been significant work done in the development of artificial

diamond detectors for radiation dosimetry. These artificial diamonds are grown by a

process of chemical vapour deposition (CVD) and they have been developed by a

number of groups 15-18. The study by Ciancaglioni et al showed that their CVD

diamond detector gave a good agreement to within 1% for measured depth doses with

field sizes down to 1×1 cm2 as compared to ionisation chamber measurements.18

Similar results were obtained in the study by Betzel et al for depth doses and relative

output factors with field sizes down to 3×3 cm2 for their CVD diamond detector.15

More recently, an artificial diamond detector has become available commercially

which has the potential for use with small field dosimetry, the PTW 60019

microDiamond detector (PTW, PTW-Freiburg, Germany).

In 2008, a new formalism for small field dosimetry was introduced by Alfonso

et al which aimed to formalize the use of Monte Carlo calculations in small field x-ray

dosimetry.4 The proposal was to introduce a field factor, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , that converts

absorbed dose to water, 𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 , for a machine-specific reference field (fmsr), with a

beam quality Qmsr, to the absorbed dose to water for the clinical field size of interest

(fclin) of beam quality Qclin . This can be mathematically expressed as:


detector 29

𝐷𝑤,𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛 = 𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 ∙ Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


Alfonso et al noted that the field factor, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , can be calculated directly as

a ratio of absorbed doses to water using Monte Carlo simulations alone or can be

measured as a ratio of detector readings multiplied by a Monte Carlo calculated


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 .

In the present work, we evaluate a newly released synthetic diamond detector,

the PTW 60019 microDiamond, for small field size x-ray beam dosimetry. Reference

dosimetry data used to compare the microDiamond detector were taken with a IBA

SFD and a PTW 60012 E diode. These diodes were recently used by Chalkley et al19

to compare with the microDiamond detector for a CyberKnife system. Monte Carlo

methods were used to calculate field factors, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for a Novalis Trilogy linear

accelerator equipped with circular cones in the range of 4-30 mm diameter. From these

field factors, we have derived the correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for the new PTW

60019 microDiamond detector for 6 MV stereotactic radiosurgery x-ray beam.

2.2 METHODS AND MATERIALS

The 6 MV SRS x-ray beam used in this work was produced by a Novalis Trilogy

linear accelerator (Varian Medical Systems, Palo Alto, USA). This beam uses a thin

flattening filter in order to produce a higher dose rate of up to 1000 MU per minute.20-

22 Beam collimation for the SRS x-ray beams was achieved by using the BrainLab

circular cones (BrainLab, Germany) of 4, 7.5, 10, 20 and 30 mm diameter as defined

at the isocentre. The X and Y collimator jaws were set to 5 cm for all measurements

with these circular cones.

The PTW 60019 microDiamond detector was compared with the PTW 60012 E

diode detector (PTW, PTW-Freiburg, Germany) and an IBA SFD solid state diode

(IBA, Schwarzenbruck, Germany). Relative dosimetry data were collected consisting

of percentage depth doses (PDDs) and cross profiles measured for the SRS circular

cones. All measurements were acquired in a large scanning PTW MP3 water phantom

(PTW, Freiburg, Germany) at an SSD of 100 cm. For the depth dose measurements,

we used a step size of 1 mm for the first 20 mm from the surface and a step size of 2.5

mm for greater depths.


detector

For all measurements with the IBA SFD diode and the PTW 60012 E diode the

detectors were oriented parallel to the central axis of the x-ray beam. Similarly,

percentage depth dose and field factor measurements with the PTW 60019

microDiamond detector were acquired with the detector oriented parallel to the x-ray

beam as per manufacturer recommendations. For measurements of cross profiles with

the PTW 60019 microDiamond detector, one set of measurements was obtained with

the detector oriented parallel to the central axis of the beam, and another obtained with

the perpendicular orientation.

Field factors were measured with the IBA SFD and PTW 60012 E. The field

factors were derived by using the daisy-chaining approach outlined by Dietrich et al23.

These field factors were used as the reference values to compare with the values

measured by PTW 60019 microDiamond detector. A previously verified and published

Monte Carlo model using BEAMnrc for a Novalis linear accelerator equipped with

circular cones was used to calculate field factors, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for cone diameters in

the range of 4 to 30 mm.24 In this model, the DOSXYZnrc user-code (V4 r2-3-0) was

used to calculate these field factors in water. Voxel sizes of 0.250.250.25mm3 were

used to score the dose. To model electron transport as accurately as possible, a global

ECUT of 0.521 MeV was specified and the EXACT boundary crossing algorithm was

turned on for the dose calculations.25-27 We then used these Monte Carlo calculated

field factors, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , to determine the correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , for the PTW

60019 microDiamond detector.

2.3 RESULTS

2.3.1 Percentage depth doses

Figure 2-1 shows the depth doses measured with the PTW 60019

microDiamond, the PTW 60012 E and the IBA SFD diode detector for the 4, 7.5, 10

and 30 mm circular cones. The agreement in depth doses between the two detectors

for all the field sizes studied was generally better than 1% with a maximum difference

of 1.5%. This level of agreement is consistent with the results of Ciancaglioni et al

who found differences of up to 2% for their CVD depth doses of a 1×1 cm2 10 MV x-

ray beam which were compared to a PTW PinPoint ionization chamber18.

It should be noted that for the depth dose measurements, no corrections were

made for dose rate response variations, such as those that have been applied for dose


detector 31

measurements often performed when using natural diamond detectors. In addition, no

corrections have been made in terms of the ratio of the stopping power of the PTW

60019 microDiamond detector and the stopping power of water. Both the PTW 60012

E and the IBA SFD diode detectors were tested for dose rate dependence by measuring

a PDD in a 10×10 cm2 field size and compared to a PDD measured with an ionisation

chamber. All PDDs were within 0.5% of each other at all depths. This confirms that

the diodes were not dependent on dose rate.

Figure 2-1 Percentage depth doses measured with a PTW 60019 microDiamond

detector (O), IBA SFD Diode (X) and PTW 60012 E Diode detector (+) for 4, 7.5, 10

and 30 mm circular cones at SSD 100 cm.


detector

2.3.2 Cross profiles and penumbra

Cross profiles measured with the PTW 60019 microDiamond and the IBA SFD

diode for the 4, 7.5, 10 and 30 mm circular cones at a depth of 10 cm are shown in

Figure 2-2 and beam profile parameters are presented in Table 2-1. Note that in Figure

2-2 only half profiles are presented to highlight the penumbral effects for the three

cases. For comparison, the IBA SFD diode was chosen over the PTW 60012 E diode

for these measurements due to its small diameter which gives a superior spatial

resolution by minimizing volume averaging effects across the penumbra.

The influence of the orientation of the PTW 60019 microDiamond detector is

most apparent in the data shown in Table 2-1; with the detector oriented perpendicular

to the beam central axis the penumbrae are consistently smaller than the IBA SFD

diode, whereas with parallel orientation the penumbrae are broader. This is attributed

to the cross sectional area of the detector causing volume/area averaging during the

measurements, with the IBA SFD diode being 0.6 mm in diameter and the PTW 60019

microDiamond detector being 2.2 mm in diameter for parallel orientation and 1 m

thickness for perpendicular orientation. Qualitatively, this is most apparent in the 4

mm cone profiles as shown in Figure 2-2.


detector 33

Figure 2-2 Half cross profiles measured at a depth of 10.0 cm with a PTW 60019

microDiamond detector in parallel orientation (), perpendicular orientation (), and

IBA SFD diode (X) for 4, 7.5, 10 and 30 mm circular cone at SSD of 100 cm.

Table 2-1 Penumbra (80%-20%) and FWHM measurements by an IBA SFD diode

and a PTW 60019 microDiamond detector

Cone

Diameter

(mm)

Penumbra FWHM

IBA

SFD

microDiamond (mm) IBA

SFD

microDiamond (mm)

(mm) Parallel Perpendicular (mm) Parallel Perpendicular

4 1.2 1.7 1.1 4.3 4.3 4.1

7.5 1.5 2.0 1.4 8.1 8.0 8.0

10 1.7 2.3 1.5 11.0 10.9 10.8

30 2.4 2.7 2.3 32.3 32.2 32.2


detector

2.3.3 Field factors, Ω𝐐𝐜𝐥𝐢𝐧 ,𝐐𝐦𝐬𝐫

𝐟𝐜𝐥𝐢𝐧,𝐟𝐦𝐬𝐫 , and correction factors 𝒌𝑸𝒄𝒍𝒊𝒏,𝑸𝒎𝒔𝒓

𝒇𝒄𝒍𝒊𝒏,𝒇𝒎𝒔𝒓

Table 2-2 also shows the field factors measured with PTW 60012 E, IBA SFD

and PTW 60016 microDiamond detectors. Monte Carlo calculated field factors,


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , are also shown in this table. The uncertainty in our Monte Carlo

simulations was within 0.5%. The type A uncertainty for our measurements was

estimated to be within 0.5% (1 SD).

Table 2-3 shows the corrections factor, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , derived from the Monte Carlo

field factors calculated in water and measurements in water for the PTW 60016

microDiamond detector for a Novalis Tx equipped with circular cones and using a 6

MV SRS x-ray beam. Please note that the Monte Carlo simulations were simulations

of Dose to Water and did not include the geometric details of the detector construction

which were intellectual property of the vendor at the time of Monte Carlo simulations.

Table 2-2 Field factors obtained using a PTW 60012 E diode, an IBA SFD diode and

PTW 60019 microDiamond detector at a depth of 1.4 cm for a 6 MV SRS x-ray beam

on a Novalis Tx equipped with circular cones at an SSD of 100 cm. The uncertainties

were up to 0.5% (1 SD) for all detectors.

Cone

diameter

(mm)

Depth

(cm)

PTW

60012 E

IBA

SFD

PTW 60019

microDiamond

Monte Carlo

relative output

factor,



30 1.4 0.940 0.943 0.944 0.959

20 1.4 0.927 0.925 0.929 0.955

10 1.4 0.860 0.851 0.856 0.870

7.5 1.4 0.808 0.798 0.799 0.811

4 1.4 0.664 0.662 0.644 0.649


detector 35

Table 2-3 Monte Carlo calculated correction factors for PTW 60019 microDiamond

detector at a depth of 1.4 cm for a Novalis Tx equipped with circular cones using a 6

MV SRS x-ray beam

Cone diameter

(mm)

Correction factor,

𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


30 1.016

20 1.027

10 1.015

7.5 1.013

4 1.006

2.4 DISCUSSION

As mentioned in the results section, the IBA SFD detector has superior spatial

resolution due to its smaller physical size. Therefore, it is important to note that

although there was a variation in the penumbra results measured in both the parallel

and vertical orientation, the actual values measured for the FWHM did not vary much

regardless of the detector orientation. The largest FWHM deviation noted in Table 2-

1 was 0.2 mm for the 4 and 10 mm cones. A similar trend was reported by Monasor et

al28 where they studied the performance of several detectors for measurements of small

field sizes down to 6 x 6 mm2. The detectors in their work included the PTW 60019

microDiamond detector as well as the IBA SFD detector. Their measurements were

performed in the parallel orientation. The results for these two detectors showed a

maximum deviation of 0.24 mm for the field size or FWHM and a maximum deviation

of 0.25 mm for the penumbra measurements. This work is in agreement with the results

presented above.

The field factors shown in Table 2-2 are within 2.7% to those published by

Garcia et al which included BEAMnrc Monte Carlo calculations and Gafchromic

EBT2 measurements29. However, there was a very close agreement in the relative

output factor for the 4 mm cone to within 0.2% as compared to this work. The

difference in Monte Carlo derived field factors can be attributed to parameterization

of the head component in the Monte Carlo model used. In addition, the selection of the


detector

energy of the incident electron beam onto the target as well as spot size distribution

has been shown to affect output correction factors30-32. Therefore, we expect that there

will be differences in field factors due to the uncertainties in the measurements and

Monte Carlo calculations on linear accelerators even between studies that used the

same model of linear accelerator.

Bassinet et al derived the output field factors from passive detector

measurements being Gafchromic EBT2 film and LiF TLDs and subsequently derived

a field factor from the mean doses from both detectors7. Our results differ from those

of Bassinet et al by up to 3.7% which is attributed to several factors. First, the work by

Bassinet et al was performed on a Varian Clinac accelerator using a standard 6 MV x-

ray beam. In comparison, the present work was performed on a Novalis Trilogy with

a 6 MV SRS beam which has a special flattening filter to produce higher dose rate x-

ray beams for SRS treatments. This difference can contribute to a different spectrum

and different output even for linear accelerators with a similar head geometry.

Additionally, the present work utilized a 5 × 5 cm2 jaw size for all measurements and

simulations where Bassinet et al varied their jaw size with differing cones7.

A recent paper by Chalkley et al19 demonstrated that the new PTW 60019

microDiamond detector has an excellent spatial resolution, dose-rate independence

and water equivalence for small fields ranging from 5 to 60 mm in diameter and for a

CyberKnife system. Those findings agree with the present work with the experimental

exception that we used a Novalis Trilogy linear accelerator. Our results also show

minimal dose rate dependence when compared to the PDDs measured by the IB SFD

and PTW 60012 E detectors. They found that for the 5 mm collimator, the

microDiamond is within 1% of the Monte Carlo corrected values, compared with the

5% and 10% correction factors for the diodes and ionization chambers, respectively19.

According to the Alfonso et al4 formalism the correction factors,


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , are machine specific which in this case it is a Novalis Trilogy equipped

with circular cones. Therefore, it should be noted that these correction factors only

apply to the cones and at an SSD of 100 cm.

Furthermore, the it should be noted that the correction factors presented in Table

2-3 are machine-specific. In this work, the linear accelerator used is a Novalis Trilogy

using the special 6 MV SRS X-ray beam mode, equipped with Brainlab circular cones


detector 37

and with a field size set to 5 × 5 cm2. This is in contrast to the IAEA TRS-483 Code

of Practice where field sizes produced by tertiary collimators are grouped into cone-

based and multi-leaf based as a single group.

2.5 CONCLUSION

In this work, we have evaluated the PTW 60019 microDiamond detector for the

dosimetry of small x-ray fields as used in stereotactic radiosurgery. This synthetic

diamond detector has been shown to possess good dosimetric properties for depth

doses, profiles and field factor measurements in the fields studied. The correction

factors supplied in this study apply for use in a Novalis Trilogy linear accelerator

equipped with Brainlab circular cones and in a 6 MV SRS x-ray beam. For cross

profile measurements, sharper penumbra measurements can be obtained with the

detector oriented perpendicular to the beam central axis.

2.6 ACKNOWLEDGEMENTS

The authors would like to acknowledge Dr Tanya Kairn for useful discussion

and comments to the work presented in this paper. Dr Scott Crowe was funded by

Australian Research Council project LP110100401. Computational resources and

services used in this work were provided by the High Performance Computing and

Research Support Unit, Queensland University of Technology (QUT), Brisbane,

Australia. We would also like to thank Brainlab for providing the specifications of the

circular cones.

2.7 CONFLICT OF INTEREST

The authors declare no conflict of interest

2.8 REFERENCES

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41




Chapter 3: Monte Carlo calculated output correction factors for Gafchromic EBT3 film for dosimetry in

stereotactic radiosurgery ray fields 43

Chapter 3: Monte Carlo calculated output

correction factors for

Gafchromic EBT3 film for

dosimetry in stereotactic

radiosurgery ray fields

Overview

There is a wide range of published literature on the use of radiochromic film for

measurements in small x-ray fields. However, it is interesting to note that most

publications have assumed that radiochromic film is water equivalent and would not

need correction for small field output factor measurements. The recently released

IAEA TRS-483 Code of Practice assumes that radiochromic film is a correction-less

detector and it was even used as part of the set of data to derive field output correction

factors. However, to date there is limited literature showing how correction-less

radiochromic film really is. The work published in this chapter discusses the water

equivalence of radiochromic film through Monte Carlo simulations. It follows from

the previous chapter on finding new knowledge and techniques that are useful for small

field dosimetry and it employs the exact geometric and material parameters of

Gafchromic EBT3 film in water using the free software programs BEAMnrc and

DOSXYZnrc to simulate radiation transport. This publication follows in the general

theme of this thesis which involves Brainlab circular cones down to 4 mm in diameter.

44 Chapter 3: Monte Carlo calculated output correction factors for Gafchromic EBT3 film for dosimetry in

stereotactic radiosurgery ray fields


Title: Monte Carlo calculated output factors for Gafchromic EBT3 film for

dosimetry in small stereotactic radiosurgery fields

Authors: Johnny E Morales, Martin Butson, Robin Hill, Scott B Crowe, Jamie

V. Trapp


Performed all Monte Carlo calculations. Involved in the project design and wrote

the entire manuscript.

Martin Butson

Provided advice as required. Helped with interpretation of results. Provided

feedback on manuscript write up.

Robin Hill



Scott B. Crowe



Jamie Trapp


results. Edited manuscript and contributed to the write up.

Journal: Australasian Physical and Engineering Sciences in Medicine

Status: Submitted 2019 – Under Review



ABSTRACT

Purpose: To calculate small field output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


Gafchromic EBT3 film using Monte Carlo simulations. These factors were determined

for a Novalis Trilogy linear accelerator equipped with Brainlab circular cones with

diameters of 4.0 to 30.0 mm.

Methods: The BEAMnrc Monte Carlo code was used to simulate the Novalis

Trilogy linear accelerator and the Brainlab cones with diameters 4.0 to 30 mm. The

DOSXYZnrc code was used to simulate Gafchromic EBT3 film with the atomic

composition specified by the manufacturer. Small field correction factors were

calculated according to new IAEA TRS-483 Code of Practice for small field

dosimetry. The depth of calculation was 10 cm and a source to surface distance of 100

cm. The x-ray beam used in the simulations was a 6 MV SRS.

Results: The Monte Carlo calculated output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


EBT3 Gafchromic film ranged between 0.998 to 1.004 for Brainlab circular cones with

diameters between 5.0 and 30.0 mm. For a diameter less than 5 mm the Monte Carlo

calculated output correction factor was 0.992.

Conclusions: For field sizes above 5 mm diameter, EBT3 Gafchromic film can

be considered to be correction less. For field sizes below 5 mm diameter Gafchromic

EBT3 film needs a small correction of 0.992.



3.1 INTRODUCTION

Small field dosimetry has a number of challenges which need to be resolved in

order to provide accurate radiation dosimetry1. Some of the limitations include the

issue of a lack of charged particle equilibrium, occlusion of the x-ray source, under-

response or over-response of the detector material relative to water and volume

averaging effects due to the size of the detector relative to the radiation beam size1-3.

To assist the medical physics community, the IAEA has recently released the IAEA

TRS-4834 code of practice (COP) which provide guidance on measuring small field

dosimetry data including field output correction factors. In order to derive small field

output correction factors, the TRS-483 COP used three types of data sets: i) reference

detectors which were perturbation free except for volume averaging, ii) reference

detectors with known output correction factors and iii) Monte Carlo calculated output

correction factors.

In the first case, where reference detectors were perturbation free except for

volume averaging, the published experimental data were obtained by comparing the

field size dependence of the small field detector with that of another small field

reference detector. In this scenario, the reference detector was assumed to be

perturbation free except for volume averaging. According to the TRS-483 COP, this

is the case for reference detectors with radiological properties and densities that are

similar to those for water. The Code of Practice assumes that because they are a ratio

in equation (47)4 (see page 121 in COP4) then no correction is required other than

volume averaging. Equation (47)4 in the COP was written as follows:


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 = 𝑀𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛

𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 ∙ 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


Where, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 is the field factor, for clinical field fclin, relative to the machine

specific field fmsr of quality Qmsr, 𝑀𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛 and 𝑀𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 are the readings for clinical field

and the machine specific reference field. 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 is the field output correction factor.

The TRS-483 COP states that such reference detectors can be Alanine, TLDs,

organic scintillators and radiochromic film. For these detectors the correction factors ,


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑], were derived with equation (59)4:




𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑] =𝑀𝑄𝐶𝑙𝑖𝑛

[𝑟𝑒𝑓]×𝑘𝑣𝑜𝑙/𝑀𝑄𝑚𝑠𝑟[𝑟𝑒𝑓]

𝑀𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]/𝑀𝑄𝑚𝑠𝑟

[𝑠𝑓𝑑]

In the second case, for reference detectors with known output correction factors,

the Code of Practice chose correction factors from independent data sets which

differed by no more than 5%. The correction factors derived from experimental data

were then calculated using equation (60)4:


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑] =𝑀𝑄𝐶𝑙𝑖𝑛

[𝑟𝑒𝑓]×𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑟𝑒𝑓]/𝑀𝑄𝑚𝑠𝑟[𝑟𝑒𝑓]

𝑀𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]/𝑀𝑄𝑚𝑠𝑟

[𝑠𝑓𝑑]

Some of the examples of these detectors included unshielded diodes, stereotactic

diodes, natural and artificially grown diamond detectors and liquid ionisation

chambers all of which are readily available. The COP provides detector specific field

output correction factors for several commercially available detectors in tables 23 to

27 in Section 6.

In the third case, for Monte Carlo calculated output correction factors, the


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑] factors were derived according to the international formalism by

Alfonso et al5 (see equation 8) and according to equation (61)4 in the COP:


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑] =𝐷𝑤,𝑄𝐶𝑙𝑖𝑛

/𝐷𝑤,𝑄𝑚𝑠𝑟

𝐷𝑑𝑒𝑡,𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]/𝐷𝑑𝑒𝑡,𝑄𝑚𝑠𝑟

[𝑠𝑓𝑑]

In this equation Dw,Q stands for the average absorbed dose to water scored in a

small voxel at the reference point in homogeneous water in a field of quality Q and

Ddet,Q [sfd] stands for the average dose scored in the small field detector [sfd] in a field

of quality Q. The TRS-483 COP excluded data obtained through a “hybrid

procedure.” This hybrid procedure was considered to be one where Monte Carlo

calculated output factors in water were combined with measured ratios of detector

readings6. The TRS-483 COP stated that even for the best commissioned Monte Carlo

model, one cannot assume that the simulation and the measurements corresponded to

the same particle fluence distribution.

It must be noted that radiochromic film, according to the TRS-483 COP, is to be

classified as a reference detector for small field dosimetry. However, to date, there has



been no Monte Carlo calculated output factors for radiochromic film in the context of

the formalism for small field dosimetry presented in TRS-483, more specifically, using

equation (61). An earlier publication by Sutherland et al7 presented Monte Carlo

calculations for the earlier version of radiochromic Gafchromic EBT and EBT2 film.

The emphasis in that work was on the energy-dependence of radiochromic film as

opposed to deriving correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑], for small field dosimetry. In

addition, the EBT and EBT2 versions of film have been discontinued and are no longer

available.

In this work, we performed Monte Carlo simulations of Gafchromic EBT3 film

in water to determine output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑], for equation (61)4 in

the COP, for small circular fields produced by a Novalis Trilogy linear accelerator

(Varian Medical System, Palo Alto, USA). The circular fields were produced by

Brainlab circular cones with diameters from 4 to 30 mm. The megavoltage x-ray beam

was 6 MV SRS.

3.2 MATERIALS AND METHODS

Monte Carlo calculations for dose to detector, being Gafchromic EBT3 film,

were performed using the atomic composition, geometry and dimensions as specified

by the manufacturer (Ashland ISP Advanced Materials, NJ, USA). Gafchromic EBT3

film consists of a matte Polyester layer 125 m thick, an active layer of 28 m

thickness and a second Matte Polyester layer 125 m thick. The atomic composition

by weight for Polyester was H (4.2%), C (62.5%) and O (33.3%). The atomic

composition by weight for the Active layer was H (8.8%), Li (0.6%), C (51.1%), O

(32.8%) and Al (6.7%). Dose to detector was scored using voxel size of 0.5 × 0.5 ×

0.028 mm3. For the term in the denominator the calculations performed were for dose

to the EBT3 active layer using a voxel size of 0.5 × 0.5 × 0.028 mm3. The Monte Carlo

calculations were performed at a depth and 10.0 cm in water. Dose to water

calculations were reproduced using exactly the same voxel geometry and size and

depth.

The Monte Carlo code used for these calculations was DOSXYZnrc 8 user code

V4 2.3.3 (NRC, Ottawa, Canada) in conjunction with a BEAMnrc model 9. The

corresponding atomic composition for each layer of the EBT3 film was not part of the



standard DOSXYZnrc package therefore a new PEGS4 data file was created, using the

EGSnrcMP 10 code, which contained the specific components.

The following parameters in DOSXYZnrc were used for all calculations: The

energy cut-off parameters for electrons were AE= 0.512 MeV and UE=25 MeV and

for photons the energy cut-off parameters were AP=0.01 MeV and UP = 25 MeV. The

global electron cut-off energy, ECUT, was set to 0.521 MeV and the global photon

cut-off energy, PCUT, was set to 0.01 MeV. A global ECUT of 0.521 MeV

corresponded to an electron range of approximately 2.5 m 11 in water which was less

than one third of the film active layer thickness. This was selected as a rule of thumb

so that the ECUT value be chosen corresponded to an electron range which was less

than one third of the smallest dimension in a dose scoring region which in this case

was 28 m. The HOWFARLESS option was switched off because this option is only

recommended for use in homogeneous media8 . If it was left turned on, then the voxel

boundaries would be ignored and the charged particles would take longer steps. The

boundary crossing algorithm used was EXACT with the electron-step algorithm set to

PRESTA-II 11, 12.

3.3 RESULTS

Table 1 presents the results obtained for 𝐷𝑤,𝑄𝐶𝑙𝑖𝑛/𝐷𝑤,𝑄𝑚𝑠𝑟

and 𝐷𝑑𝑒𝑡,𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]/

𝐷𝑑𝑒𝑡,𝑄𝑚𝑠𝑟[𝑠𝑓𝑑] from Monte Carlo simulations . Column 1 in the table shows the

diameter for each cone simulated. The second column shows the depth at which the

simulation was performed, in this case it was 10.0 cm for all cones. This is the refence

depth specified in the COP4. Column 3 shows the values obtained for 𝐷𝑤,𝑄𝐶𝑙𝑖𝑛/𝐷𝑤,𝑄𝑚𝑠𝑟

and column 4 shows the values obtained for 𝐷𝑑𝑒𝑡,𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]/𝐷𝑑𝑒𝑡,𝑄𝑚𝑠𝑟

[𝑠𝑓𝑑]. These

values were then used to derive detector correction factors 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑] for

Gafchromic EBT3 film were derived according to equation (61) in the COP. All results

were calculated at 100 cm SSD as specified in the COP.



Table 3-1 Monte Carlo calculated correction factors, kQclin,Qmsr

fclin,fmsr [sfd], for Gafchromic

EBT3 film for a 6 MV SRS beam Cone

diameter

(cm)

Depth

(cm) 𝐷𝑤,𝑄𝐶𝑙𝑖𝑛

/𝐷𝑤,𝑄𝑚𝑠𝑟

𝐷𝑑𝑒𝑡,𝑄𝐶𝑙𝑖𝑛[𝑠𝑓𝑑]

/𝐷𝑑𝑒𝑡,𝑄𝑚𝑠𝑟[𝑠𝑓𝑑]


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑]

4.0 10.0 0.356 0.359 0.992

7.5 10.0 0.452 0.452 1.000

10.0 10.0 0.488 0.489 0.998

20.0 10.0 0.551 0.551 1.000

30.0 10.0 0.570 0.568 1.004

3.4 DISCUSSION

The results show that the Monte Carlo calculated correction factors,


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑], are within 0.4 % for diameters values greater than 5 mm for the

Brainlab circular cones. The COP provides correction factors in Tables 23 to 27 (see

pages 124 to 136) for several detectors including ionization chambers, shielded and

unshielded diodes, diamonds, synthetic diamond, plastic organic scintillator to name a

few. However, it does not provide correction factors in these tables for Alanine, TLDs

and radiochromic film on the basis that they are regarded as reference detectors and

they only need correction for volume averaging effects. As mentioned in the

introduction, for radiochromic film, this assumption has been implied due to its high

spatial resolution and radiological water equivalence. But to date, there has been no

definitive proof to show that this is case using a full model of the film and water for

small fields. Some of the reasons given are that no matter how good a Monte Carlo

model is, it will not provide the exact same fluence and beam spectra as that of a linear

accelerator4. In comparison, the COP only provides the 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 correction factors

for a few nominal x-ray beam energies and are determined across a range of

radiotherapy equipment with slightly different beam spectra.

In this work, we have shown by simulating the conditions for a similar particle

fluence in the clinical field, fclin, and the machine specific reference field, fmsr, as well

as using the correct atomic composition of the EBT3 film, the correction factors,


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , are close to unity. Therefore, our results are complementary to the results

in the COP and other studies of small field dosimetry2, 13-15.



The work by Underwood16 et al assumed that the correction factors for EBT3

film were unity when used to determine relative output factors for a number of

prototype diode detectors. Wegener et al [2017] used Gafchromic EBT3 film for

relative dosimetry measurements of small fields down to 5 mm field size and for

comparison with a variety of solid state detectors. Our results confirm the validity of

those assumptions and the subsequent correction factors.

In a more extensive study, Yarahmadi17 et al determined output factors for small

x-ray fields. Firstly, they measured using an SFD diode and Gafchromic EBT3 film

for field sizes down to a nominal field size of 6 mm for an Elekta 6 MV x-ray beam.

In addition, they performed Monte Carlo calculations of the dose to the active volume

of the SFD as well as to a small voxel of water. The overall agreement was good

between all output factors with a maximum difference of 2% at 8 mm but only a

difference 0.5% for the 6 mm field size. While they have not directly modelled the

EBT3 film in the Monte Carlo calculations as has been done in this work, our work is

very consistent with their results. This provides confirmation that the 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 [𝑠𝑓𝑑]

correction factors are close to unity.

There is no doubt that it is quite challenging to perform dosimetry measurements

using radiochromic film. The final outcome can be influence by several factors

including: scanning orientation dependence, energy dependence, post coloration

behaviour, temperature dependence, ambient light sensitivity and absorption spectra18-

26. However, as specified in the COP, if an accurate and reproducible calibration and

readout procedure is consistently followed in the local department, Gafchromic EBT3

film can be used for determination or verification of small field output factors down to

5 mm without the need for correction.

The COP specifies that the only correction needed for radiochromic film, TLDs

and Alanine would be the volume average effect. Some studies have shown that

radiochromic film itself can help to quantify the effect of volume averaging2 in small

fields and that the volume average effect could be removed altogether from

radiochromic film measurements27. Thus, our study further supports the assumption

that radiochromic film is a correctionless dosimeter and that if used appropriately it

can be used as a reference detector for small field dosimetry.



Monte Carlo methods used to simulate detector response can contain limitations

with regard to detector geometry. This has been a tremendous challenge in recent years

as the geometry of a detector component could potentially affect the overall Monte

Carlo simulation. This can be the case for modelling semi conductor detector such as

diodes, diamonds and synthetic diamonds. Radiochromic film can also have its

limitations as radiochromic film sensitive material can only be model as a single layer

of a particular material. While this is true in reality radiochromic film is made up of

micro granules surrounded by dye material. The granules can have random orientations

and it is best to simulate the active layers as a whole. These factors can add to the

uncertainty.

3.5 CONCLUSION

Our study shows that the relative output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟


Gafchromic EBT3 radiochromic film for a 6 MV SRS x-ray beam are within 0.4% for

circular fields from 5 to 30 mm in diameter at an SSD of 100 cm and 10.0 cm depth.

The correction factors were calculated using equation (61)4 in the recently released the

IAEA TRS-4834 code of practice (COP). The study shows that for smaller field sizes

below 5 mm a correction close to 1% could be needed. However, this uncertainty

would need to be incorporated in the overall uncertainty budget for the readout system

used during calibration and readout procedures for Gafchromic EBT3 film.


Some of the calculations were performed on an in-house computer cluster

comprising of 64 computing nodes connected in parallel and running on Linux CentOS

version 6. This cluster was constructed from funds provided by QUT. Also, some of

the computational resources and services used in this work were provided by the HPC

and Research Support Group, Queensland University of Technology, Brisbane,

Australia. We would like to thank David F. Lewis of Ashland Inc. for providing the

information on the atomic composition of Gafchromic EBT3 radiochromic film

modelled in this study. We thank Brainlab for providing the specifications of the

circular cones and useful discussions.




The authors declare no conflict of interest.

3.8 REFERENCES

1 I.J. Das, G.X. Ding, A. Ahnesjo, "Small fields: Nonequilibrium radiation

dosimetry," Medical Physics 35, 206-215 (2008). 2 G. Azangwe, P. Grochowska, D. Georg, J. Izewska, J. Hopfgartner, W.

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Fukumura, K. Yajima, C. Gouldstone, P. Sharpe, A. Meghzifene, H. Palmans,

"Detector to detector corrections: A comprehensive experimental study of

detector specific correction factors for beam output measurements for small

radiotherapy beams," Medical Physics 41 (2014). 3 P.H. Charles, G. Cranmer-Sargison, D.I. Thwaites, S.B. Crowe, T. Kairn, R.T.



measurements," Medical Physics 41, - (2014). 4 I.A.E. AGENCY, Dosimetry of Small Static Fields Used in External Beam


(2017). 5 R. Alfonso, P. Andreo, R. Capote, M.S. Huq, W. Kilby, P. Kjäll, T. Mackie,


of small and nonstandard fields," Medical Physics 35, 5179-5186 (2008). 6 C. Martens, "The value of the pinpoint ion chamber for characterization of

small field segments used in intensity-modulated radiotherapy," Physics in

Medicine and Biology 45, 2519-2530 (2000). 7 J.G.H. Sutherland, D.W.O. Rogers, "Monte Carlo calculated absorbed-dose

energy dependence of EBT and EBT2 film," Medical Physics 37, 1110-1116

(2010). 8 B.R.B. Walters, D.W.O. Rogers, "DOSXYZnrc Users Manual," NRC Report

PIRS 794 (rev B)2004). 9 J.E. Morales, R. Hill, S.B. Crowe, T. Kairn, J.V. Trapp, "A comparison of



Sciences in Medicine 37, 303-309 (2014). 10 I. Kawrakow, E. Mainegra-Hing, D.W.O. Rogers, F. Tessier, B.R.B. Walters,

"The EGSnrc Code System: Monte Carlo Simulation of Electron and Photon

Transport," NRC Report PIRS 701 (2011). 11 J.H. Kim, R. Hill, Z. Kuncic, "An evaluation of calculation parameters in the


calculation," Physics in Medicine and Biology 57, N267-N278 (2012). 12 B.R.B. Walters, I. Kawrakow, "Technical note: Overprediction of dose with


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14 C. Moignier, C. Huet, L. Makovicka, "Determination of the

K<SUP>f<INF>clin,fmsr</INF></SUP>

<INF>Q<INF>clin,Qmsr</INF></INF> correction factors for detectors used

with an 800 MU/min CyberKnife (R) system equipped with fixed collimators

and a study of detector response to small photon beams using a Monte Carlo

method," Medical Physics 41 (2014). 15 F.I.O. Alexander, J. Wassim, "Correction factors for diode and diamond

detectors in the measurement of small field output factors, using film dosimetry

as reference," Biomedical Physics & Engineering Express 4, 055011 (2018). 16 T.S. Underwood, B.C. Rowland, R. Ferrand, L. Vieillevigne, "Application of

the Exradin W1 scintillator to determine Ediode 60017 and microDiamond

60019 correction factors for relative dosimetry within small MV and FFF

fields," Phys Med Biol 60, 6669-6683 (2015). 17 M. Yarahmadi, S. Wegener, O.A. Sauer, "Energy and field size dependence of

a silicon diode designed for small-field dosimetry," Med Phys 44, 1958-1964

(2017). 18 B.D. Lynch, J. Kozelka, M.K. Ranade, J.G. Li, W.E. Simon, J.F. Dempsey,

"Important considerations for radiochromic film dosimetry with flatbed CCD

scanners and EBT GAFCHROMIC film," Med Phys 33, 4551-4556 (2006). 19 A. Rink, D.F. Lewis, S. Varma, I.A. Vitkin, D.A. Jaffray, "Temperature and

hydration effects on absorbance spectra and radiation sensitivity of a

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EBT radiochromic film," Phys Med Biol 50, N281-285 (2005). 21 B. Arjomandy, R. Tailor, A. Anand, N. Sahoo, M. Gillin, K. Prado, M. Vicic,

"Energy dependence and dose response of Gafchromic EBT2 film over a wide

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factor measurements in small x-ray fields," Med Phys 43, 4687 (2016).

55




Chapter 4: An experimental extrapolation technique using Gafchromic EBT3 film for relative output factor

measurements in small x-ray fields 57

Chapter 4: An experimental extrapolation

technique using Gafchromic

EBT3 film for relative output

factor measurements in small x-

ray fields

Overview

Factors that can influence the accuracy of dosimetry measurements in small

fields include source occlusion, lack of charge particle equilibrium and detector

volume averaging. The TRS 483 Code of Practice presents correction factors for

several detectors for different beams and for different radiation equipment. However,

these correction factors only contain generic volume correction factors. Volume

averaging effects are due to the profile shape and the dose that is averaged over a given

area within the beam. Different methods that have been proposed to deal with volume

averaging effects in small field dosimetry including the use of film profiles to account

for the shape of the small field or by performing full geometry calculations using

Monte Carlo methods to obtain corrections factors. In this chapter, a novel

experimental extrapolation technique is introduced for Gafchromic EBT3 film

(Ashland, NJ, USA). The proposed technique can potentially eliminate the volume

averaging effect in small field dosimetry. Specifically, the paper demonstrates how

increasing the film scanning resolution can vastly improve the determination of output

factors for small fields.

58 Chapter 4: An experimental extrapolation technique using Gafchromic EBT3 film for relative output factor

measurements in small x-ray fields


Title: An experimental extrapolation technique using Gafchromic EBT3 film for

relative output factor measurements in small x-ray fields

Authors: Johnny E Morales, Martin Butson, Scott B Crowe, Robin Hill, J.V.

Trapp




Martin Butson

Provided advice and supervision as required. Helped with interpretation of

results. Provided feedback on manuscript write up.

Scott B. Crowe


Robin Hill


J.V. Trapp


results. Edited manuscript and contributed to the write up

Journal: Medical Physics

Status: Published July 2016


SCOPUS Author h-index: 6

http://dx.doi.org/10.1118/1.4958679



ABSTRACT

Purpose: An experimental extrapolation technique is presented which can be

used to determine relative output factors for very small x-ray fields using Gafchromic

EBT3 film.

Methods: Relative output factors were measured for Brainlab SRS cones

ranging in diameters from 4 to 30 mm2 on a Novalis Trilogy linear accelerator with 6

MV SRS x-rays. The relative output factor was determined from an experimental

reducing circular region of interest (ROI) extrapolation technique developed to remove

the effects of volume averaging. This was achieved by scanning EBT3 film

measurements with a high scanning resolution of 1200 DPI. From the high resolution

scans, the size of the circular regions of interest was varied to produce a plot of relative

output factors versus area of analysis. The plot was then extrapolated to zero to

determine the relative output factor corresponding to zero volume.

Results: Results have shown that for a 4 mm field size, the extrapolated relative

output factor was measured at a value of 0.651 ± 0.018 as compared to 0.639 ± 0.019

and 0.633 ± 0.021 for 0.5 mm and 1.0 mm area of analysis values, respectively. This

showed a change in relative output factor of 1.8% and 2.8% at these comparative

region of interest sizes. In comparison, the 25 mm cone had negligible differences in

the measured output factor between zero extrapolation, 0.5mm and 1.0 mm diameter

ROIs, respectively.

Conclusions: This work shows that for very small fields such as 4.0 mm cone

sizes, a measurable difference can be seen in the relative output factor based on the

circular ROI and size of the area of analysis using radiochromic film dosimetry. We

recommend to scan Gafchromic EBT3 film at a resolution of 1200 DPI for cone sizes

less than 7.5mm and to utilize an extrapolation technique for output factor

measurements in very small field dosimetry.



4.1 INTRODUCTION

Accurate small x-ray field dosimetry is crucial in the area of stereotactic

radiotherapy. There are a number of factors that can influence the accuracy of

dosimetry measurements such as source occlusion, lack of charged particle

equilibrium, non-tissue equivalence of the detector active volume and overlapping

penumbra1, 2. Therefore, the selection of a suitable detector is crucial and should have

minimal corrections for each of these factors where possible. A number of

commercially available detectors introduce volume averaging effects in the field

during measurements and a correction might be needed as shown by some studies3.

Volume averaging effects are specifically due to the profile shape of the small field

beams and the influence of maximum dose deposition that is averaged over a given

area within the small field beams3.

Some methods have been proposed to deal with volume averaging effects in

small field dosimetry including using film profiles to account for the shape of the small

field3, 4 or by performing full geometry calculations using Monte Carlo methods to

obtain correction factors5, 6. However, Monte Carlo simulations can be an arduous task

requiring a full model of the treatment machine, a process not readily available to many

radiation oncology departments5-10.

Previous studies have been performed using radiochromic film for small field

dosimetry such as the work by Lopez et al11-13. In this study, we introduce an

experimental extrapolation technique using Gafchromic EBT3 film (Ashland, NJ,

USA) which can potentially eliminate the volume effect in small field dosimetry. This

technique consists of varying the size of the region of interest within the scanned area

of the film. A method is presented which uses a zero area extrapolation technique

utilising high resolution scanning in order to determine the final relative output factors

for very small fields.


All relative output factor measurements were performed on the 6 MV SRS x-ray

beam as produced by a Novalis Trilogy linear accelerator (Varian Medical Systems,

Palo Alto, USA) which has a thin flattening filter in order to produce a dose rate of

1000 MU per minute 14-16. Beam collimation for the SRS x-ray beams was achieved

by using Brainlab circular applicators (Brainlab, Germany) with diameters between 4



and 30 mm as defined at the isocentre. The X and Y collimator jaws were set to 5 cm

for all measurements. The relative output factor is defined as the measurement for a

given cone at depth of 1.5 cm and SSD of 98.5 cm divided by the measurement for the

reference field size at a depth of 1.40 cm and SSD of 100 cm as per manufacturers

recommendations. The reference field size was 10 × 10 cm2. Monitor units delivered

to each film ranged between 180 MU and 350 MU depending on cone size in an

attempt to deliver absorbed doses close to 200 cGy to each film. Nine experimental

films were irradiated separately for each cone size evaluated from which the average

output factor and uncertainties were calculated.

4.2.1 Film dosimetry technique

Gafchromic EBT3 film (lot number 09151402) was used for all relative output

factors measurements in this work. All pieces of film were used and handled in the

process outlined in the AAPM TG-55 report17 and the Medical Radiation dosimetry

with radiochromic film report series18. It has been shown that Gafchromic EBT3 film

possesses a minimal x-ray energy dependence19, 20 and therefore should have a minimal

impact on output factor assessment at small field sizes using 6 MV x-rays. All films

were analysed using a PC desktop scanner and Image J (National Institutes of Health,

USA) software on a PC workstation at least 24 hours after irradiation to minimize

effects from post irradiation colouration21. The scanner was an Epson 10000XL dual

lens system desktop scanner (Epson, NSW, Australia) using a scanning resolution of

1200 pixels per inch. The images produced were 48 bit RGB colour images and

analysed with the red component of the signal making the final pixel density values 16

bit information22.

A control film was scanned with every experimental film in the same position

for each measurement. The resulting scans were then corrected for any inter scan

variations based on the control films result compared to the average result in a

technique similar to that used by Lewis et al 23.

The dose delivered to each experimental film was calculated by creation of a

calibration dose response curve for the Gafchromic EBT3 film using standard fields

of 10 x 10 cm at given applied dose levels. This was performed due to the known,

nonlinear relationship of net optical density to dose response of EBT3 Gafchromic

film when scanned using an Epson10000XL desktop scanner.



The net optical densities were calculated for each film piece using circular

regions of interest in ImageJ software. The circular region of interest was centred on

the cone produced radiation field. The relative output factors versus diameter size were

plotted in a curve. A best fit extrapolation technique was applied to the results to

determine the zero area output factor utilizing a 2nd order polynomial or linear

function based on cone size and data requirements.

Relative output factor measurements determined with a PTW 60019

microDiamond detector24-28 (PTW, PTW-Freiburg, Germany) have also been included

for comparison with the extrapolated Gafchromic EBT3 results. These measurements

were made under the same geometrical conditions as the measurements for

Gafchromic EBT3 film.

4.3 RESULTS

Figure 4-1 shows the measured values for relative output factor for a 4 mm cone

when different size circular regions of interest ranging from 1.8 mm to 0.1 mm were

used for the analysis in ImageJ. The figure also shows the extrapolation estimate of

the relative output factor for a zero volume or area calculation which was found to be

0.651 ± 0.018 for the 4 mm Brainlab cone. The extrapolation was performed using a

2nd order polynomial line of best fit to provide the best estimate at zero volume output

factor. The variation in measured output factor with different regions of interest areas

is expected to change due to a number of reasons including profile shape of the beam

and light scattering properties within the scanner and film. The uncertainty values

quoted in the figures and text are the standard deviation in measured results comparing

the nine experimental films assessed for each cone size measured. These values

combine both type A and type B errors associated with setup and dose delivery

uncertainty along with experimental film analysis errors combined. Results are quoted

as 1 standard deviation of the mean.

Figure 4-2 shows an example net optical density profile for one 4 mm cone as

measured by EBT3 film at a resolution of 1200 DPI which is the average of nine EBT3

film measurements. The central axis region of the profile is also shown in more detail

to highlight both the variation in net optical density and the uncertainty or noise level

with the film scan for this typical measurement. In terms of very small region of

interest analysis, the user should make multiple measurements using different films to



minimize the impact of selecting a small region of interest for analysis around either a

noise peak or trough, thus skewing the measured dose level either high or low by the

magnitude of the noise which in our case was found to be between 0.5% to 1.5%.

We used the extrapolation technique using a low order polynomial fit. Another

technique would be to use a 2 D-surface polynomial fit which can extrapolate to a zero

area over the entire film region of interest. This would be a similar process if a solid

detector with a fixed area of the interest had been used. Such detectors could have been

TLDs and OSLDs, where the active area would have been perpendicular to the beam’s

axis and therefore integrated over a 2 D region. The pros of the film extrapolation

technique are that it is simple, easy to perform with any software, for example Excel,

and provided a result in agreement with other establishes detectors. The cons are the

noise as shown in Fig 4.2. and an asymmetry in the beam may skew results and produce

a larger uncertainty in the penumbra region and thus in the whole region of interest.

Figures 4-3 and 4-4 show similar results but for a 25 mm cone size produced by

the same 6 MV SRS beam. The extrapolated relative output factor for this cone was

calculated to be 0.971 ± 0.017. The 25 mm cone however produces a negligible

difference in measured output factor when the diameter of the circular ROI analysis

ranges from approximately 1.8 mm down to 0.1 mm. This is expected as Figure 4-3

shows that a percentage dose plateau does exist at this field size. However, when the

same comparison was made for the very small 4 mm cone the variation in measured

output factor was 2.8 % . These results highlight the necessity of high spatial resolution

scanning and analysis for very small (4mm) cone fields measurements.



Figure 4-1 Measured and extrapolated output factor values for a 4mm Brainlab cone

using Gafchromic EBT3 film with varying sizes of analysis area. As the area of

analysis decreases, an increase in measured output factor occurs due to the non-

plateauing nature of the 4mm cone profile.

Figure 4-2 Measured net optical density profile for a 6MV SRS x-ray beam, 4 mm

cone taken as the average of nine EBT3 film measurements. The insert in the figure

includes details of the centre 1 mm of the profile.

0.6

0.61

0.62

0.63

0.64

0.65

0.66

0.67

0.68

0 0.5 1 1.5 2

Me

asu

red

Ou

tpu

t Fa

cto

r

Diameter of analysis area (mm)

EBT3 zero area extrapolation

4mm Cone Output Factor



Figure 4-3 Measured and extrapolated output factor values for a 25 mm diameter

Brainlab cone using Gafchromic EBT3 film with varying diameters of analysis area.

As the area of analysis decreases negligible differences in measured output factor

occurs due to the plateauing nature of the 25mm cone profile.

Figure 4-4 Measured net optical density profile for a 6MV x-ray beam, 25 mm

diameter Brainlab cone. Insert in the picture includes details of the centre 4 mm of the

profile for an example film showing the central plataeu effect at this field size.

Figure 4-5 shows the variation in measured output factor for these 2 circular

cones as a percentage decrease when each measured point is compared to the

extrapolated zero area output factor value. As can be seen, over the range up to 2 mm

diameter areas of analysis an approximate percentage difference for the 25 mm cone

is negligible, whereas for the 4 mm cone the value rises to approximately 3%. These

0.94

0.95

0.96

0.97

0.98

0.99

1

1.01

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2

Me

asu

red

Ou

tpu

t Fa

cto

r


EBT3 zero area extrapolation

25mm Cone Output Factor



results highlight the significance of the volume averaging effect and effective detector

size when dealing with very small fields like a 4 mm cone.

Figure 4-5 Percentage difference in output factor from the extrapolated zero area value

for the 4 and 25 mm Brainlab cones. Variations are seen using the 4 mm cone but

negligible differences calculated at larger cone sizes.

Table 4-1 shows the relative output factor determination when this technique is

applied to other cone sizes available. The uncertainties in the film measurements are

not quoted to improve clarity in the table, however, the authors note that they are of

similar size to those in Figures 4-1 and 4-2. Shown are the measured values at 0.5 mm

and 1 mm diameter circular ROI areas of analysis. Also shown are the percentage

differences or percentage under prediction in relative output factors which would occur

if the measurements were performed at either 0.5 mm or 1 mm effective detector

diameter.

-1

-0.5

0

0.5

1

1.5

2

2.5

3

3.5

4

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2

% D

iffe

ren

ce fr

om

ze

ro a

rea

valu

e


4mm and 25mm Cone Size

25mm Cone

4mm Cone



Table 4-1 Measured and extrapolated relative output factors for Gafchromic EBT3

film with different analysis diameters and for the PTW 60019 microDiamond

detector

Cone

size

(mm)

PTW

60019

Gafchromic EBT3 film – analysis

diameter

% difference

(Measured) 0 mm

(Extrapolated)

0.5 mm

(Measured)

1 mm

(Measured)

% diff

( 0 -

0.5

mm)

% diff

(0 – 1

mm)

4 0.661 0.651 0.639 0.633 1.8 2.8

7.5 0.821 0.810 0.804 0.803 0.7 0.9

10 0.879 0.870 0.869 0.866 0.1 0.5

12.5 0.912 0.903 0.902 0.901 0.1 0.2

15 0.932 0.932 0.932 0.931 0.0 0.1

17.5 0.946 0.947 0.948 0.947 -0.1 0.0

20 0.956 0.960 0.959 0.959 0.1 0.1

25 0.964 0.971 0.970 0.970 0.1 0.1

30 0.975 0.981 0.980 0.981 0.0 -0.1

These results highlight that even for a 0.5 mm effective detector size, the ROF

for the 4 mm cone could be under predicted by approximately 2 %. This effect is less

than 1 % for cone sizes of 7.5 mm or greater.

An analysis of effective detector size compared to scanning resolution for the

Epson 10000XL scanner and Gafchromic EBT3 film is shown in Table 4-2.

Table 4-2 Analysis of detector effective size versus pixels measured for various

resolutions

Scan resolution Size of

pixel

Equivalent

circle

Pixels per effective

detector size (circle)

Dots/inch Dots/mm Square

(mm)

Diameter

(mm)

0.1 mm 0.5 mm

75 2.95 0.339 0.382 0.07 1.71

100 3.93 0.254 0.287 0.12 3.04

150 5.91 0.169 0.191 0.27 6.84

200 7.87 0.127 0.143 0.49 12.20

300 11.81 0.085 0.096 1.10 27.40

600 23.62 0.042 0.048 4.38 109.00

1200 47.24 0.021 0.024 17.53 438.00

From these results, we can see that using a scanning resolution of 300 dpi and a

0.5 mm circular region of interest, 27 pixels would be averaged to count towards net

OD calculation. Whereas, using a scanning resolution of 1200 dpi, 438 pixels would



be used for analysis. Similarly, for a 0.1mm effective detector size, the values would

be 1 pixel and 17 pixels respectively for 300 dpi and 1200 dpi.

4.4 DISCUSSION

For small fields, which are almost entirely penumbral and thus non-flat at the

centre of the field, the selection of ROI size in planar measurements can significantly

impact the measurement result. This is illustrated in Figure 4-6, where 2 different ROIs

of the same field are selected; the average value for net OD of pixels within ROI A

will be greater than in ROI B due to the greater number of pixels in ROI B that are of

lesser value. Therefore, obtaining a series of ROIs of different diameters will enable

extrapolation to the zero volume to obtain the true output factor.

Figure 4-6 Exaggerated example of a small field. The average value of pixels in ROI

A will be greater than that of ROI B.

Table 4-2 highlights the standard scanning resolutions found on the Epson

10000XL flatbed scanner. In normal clinical operations, a scanning resolution of 75 to

300 dots per inch would normally suffice. However, when dealing with very small

fields the resolution plays a major role. In the table, the conversion to dots per

millimetre is shown along with the pixel square side size and equivalent circular

diameter per pixel. When these values are applied to an area analysis using a circular

region of interest, the number of effective pixels which lie in the area of analysis is

shown for effective detector diameters of 0.1 mm and 0.5 mm. This means that if an

equivalent detector size of 0.1 mm is required a scanning resolution of 300 dots per

inch or below will only give you one pixel or less information. At 1200 DPI resolution,



one would use approximately 17 to 18 pixels for optical density analysis. Using the

high resolution provides substantially more data for analysis and subsequently should

provide a more robust measurement to determine relative output factors and

uncertainty levels. For this reason, we recommend a scanning resolution of at least

1200 DPI to perform the extrapolation of the relative output factor measurements for

very small cones such as the 4 mm cone.

The importance of dosimetric accuracy for very small field relative output factor

lies in both planning data as well as experimental dose verification. Whilst planning

computer grid sizes are often larger than the sub millimetre measurement size in this

work, they still require an accurate relative maximum dose factor for each cone as their

defined value. For example, the Brainlab iPlan treatment system utilizes an adaptive

grid resolution down to 0.5 mm. By utilizing the extrapolation technique, as shown in

this work, the peak relative output factor for the very small 4 mm cone can be

accurately measured and given for dose calculation in the planning system.

Radiochromic film has been used more extensively for the determination of

small field size relative output factors and the data in Table 4-3 shows the different

DPI resolution used in various small field dosimetry studies29-36. This table shows that

no study has performed their final results with a resolution higher than 150 DPI and

most of these studies reported values based on a single scanning resolution. In addition,

no extrapolation technique was reported to have been used in any of these studies.

Table 4-3 The DPI scanning resolution used across studies for small field dosimetry

Study Film type Resolution DPI Year

Wilcox and Daskalov (Ref. 29) EBT 75 2007

Garcia-Garduno et al (Ref. 30) EBT 100 2010

Kairn et al (Ref. 31) EBT2 72 2011

Aland et al (Ref. 32) EBT2 75 2011

Huet et al (Ref. 33) EBT2 150 2012

Fiandra et al (Ref. 34) EBT2 and EBT3 72 2013

Huet et al (Ref. 35) EBT3 150 2014

Moignier et al (Ref. 36) EBT3 150 2014

Morales et al (Ref. 37) EBT3 150 2014

These resolutions are appropriate for cone sizes of 10 mm and above, however,

x-ray beams with dimensions less than 10 mm require greater resolution. Therefore, a

higher resolution and extrapolation technique are provided to best estimate the output

factors for these very small cone sizes.



The PTW 60019 microDiamond detector has been extensively studied for small

field dosimetry24-28, 37, 38. For the 4 mm diameter cone used in this study, the ROF

calculated using the zero area extrapolation technique was in good agreement to the

ROF measured with the microDiamond detector. As such, we believe this agreement

shows the suitability of using a zero area technique for these smaller fields in film

dosimetry.

4.5 CONCLUSIONS

The technique presented in this study is useful for analysis of Gafchromic EBT3

film results in order to eliminate volume averaging effects3, to measure very small

relative output factors, and to correct for volume averaging effects, which can be

caused by insufficient scanning resolution.

It provides enough detail and guidance to be able to reproduce Gafchromic EBT3

film measurements with high resolution. It has been shown that very small field

relative output factor measurements are dependent on the effective detector size. When

cones of size 4 mm are used, Gafchromic EBT3 film can be used to measure the

relative output factor. The measurement value is, however, affected by the average

volume or effective diameter size of the area of analysis and up to 2.8% differences

were measured when using effective detector sizes ranging from 0.0 mm (extrapolated)

to 1 mm effective diameter sizes. Field sizes ranging from 7.5 to 30 mm produced

negligible differences for the same effective detector diameters.



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75


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Chapter 5: A novel extrapolation method using OSL detectors for very small field output factor measurement for

stereotactic radiosurgery 77

Chapter 5: A novel extrapolation method using

OSL detectors for very small field

output factor measurement for

stereotactic radiosurgery

Overview

An extrapolation technique for small field measurements using film was

introduced in Chapter 4. The same technique is here applied to another type of detector

with a potential use in small field dosimetry, Optically Stimulated Luminescence

Dosimeters (OSLDs). OSLDs are sometimes used as personal dose monitors or for in

vivo dosimetry for patients undergoing radiotherapy treatment. OSLDs are very

popular because they can be read out and analysed shortly after irradiation, unlike film,

which can require a longer waiting period. In this paper, a new modified OSLD

dosimeter was constructed from a commercial NanoDot OSLD. The sensitive area of

the OSLD was reduced in size to allow an extrapolation towards zero area. This

extrapolation technique was applied for a range of field sizes from 4 to 30 mm diameter

on a 6 MV SRS beam. The application of this techniques provides another option when

determining very small field output factors.

78 Chapter 5: A novel extrapolation method using OSL detectors for very small field output factor measurement

for stereotactic radiosurgery


Title: A novel extrapolation method using OSL detectors for very small field

output factor measurement for stereotactic radiosurgery

Authors: Te-An Polly Huang, Johnny E. Morales, Ethan Butson, Annie

Johnson, Martin Butson, Robin Hill


Participated in the measurements. Involved in the project design and took part in

writing the manuscript. Corresponding author.

Te-An Polly Huang

Participated in the measurements. Involved in the project design and took part in

writing the manuscript.

Ethan Butson


Annie Johnson


Martin Butson


Robin Hill







ABSTRACT

Appropriate methods for the determination of very small x-ray beam output

factors are essential to ensure correct clinical outcomes for stereotactic radiosurgery.

To date, substantial work has been performed in identifying and quantifying suitable

dosimeters for relative output factor (ROF) measurements including recent IAEA

published recommendations. In this work, we provide a novel method using varying

sized optically stimulated luminescent dosimeters (OSLDs) to determine ROFs. This

involves applying an extrapolation technique to assess ROFs for 6MV SRS x-ray

beams with field diameters ranging from 4 to 30 mm as defined by the Brainlab SRS

cones. By combining the use of multiple sized OSLDs and water droplets to remove

air gaps located around the OSLD detectors, both volume averaging and density

variation effects were minimised to estimate ROFs for an extrapolated zero volume

detector. The measured results showed that for a 4 mm diameter cone, the ROF was

0.660 ± 0.032 (2SD) as compared to 0.661 ± 0.01 and 0.651 ± 0.018 for the PTW

600019 microDiamond detector and Gafchromic EBT3 film respectively. Whilst the

uncertainties were larger than conventional detectors, the technique shows promise

and improvements in accuracy may be obtained by higher quality manufacturing

techniques. Based on these results, using varying sized OSLDs and an extrapolation

technique shows promise for use as an independent verification tool for very small x-

ray field ROFs in the clinical department.



5.1 INTRODUCTION

The accurate determination of radiation dose is essential for patient treatments

in radiotherapy, in particular for high dose stereotactic radiosurgery 1, 2. There are a

number of key factors which influence dosimetric accuracy for small fields including

lack of charged particle equilibrium, volume averaging of the detector, occlusion of

the source, non-tissue equivalence of the detector and surrounding volume and

overlapping penumbra 1, 3, 4. For these reasons, many clinically available dosimeters

have been investigated and used for small field dosimetry each with their own

respective advantages and disadvantages. These dosimeters include small volume

ionisation chambers, diodes, synthetic and natural diamond detectors, plastic

scintillators, TLDs, alanine, MOSFETs and radiochromic film 5-15. However, many of

these detectors, due to their size and construction, require both field size and detector

density specific correction factors in order to accurately determine the dose output 16.

The formalism for small field dosimetry as introduced by Alfonso et al and used

in the IAEA TRS 483 Code of Practice provides the methodology for determining

correction factors for a wide range of detectors and geometries 17, 18. The formalism

defines the field factor, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , that converts absorbed dose to water, 𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 , for

a machine-specific reference field (fmsr), with a beam quality Qmsr, to the absorbed dose

to water for the clinical field size of interest (fclin) of beam quality Qclin . This can

mathematically expressed as:

𝐷𝑤,𝑄𝑐𝑙𝑖𝑛

𝑓𝑐𝑙𝑖𝑛 = 𝐷𝑤,𝑄𝑚𝑠𝑟

𝑓𝑚𝑠𝑟 ∙ Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 1)

Alfonso et al noted that the field factor, Ω𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , can be calculated directly as

a ratio of absorbed doses to water using Monte Carlo simulations alone or can be

measured as a ratio of detector readings multiplied by a Monte Carlo calculated


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 . For most clinics, Monte Carlo calculations are not

feasible and one relies on published data. This was the rationale for the publication of

the IAEA TRS 483 Code of Practice which has extensive tables of the factor


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 for a wide range of detectors, field sizes and beam energies.

An alternative approach has been to use a radiation dosimeter for which the


𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 is determined to be close to unity even at smaller field

sizes. This has been assumed and/or established for several detectors being TLDs,



small volume alanine detectors and radiochromic film 5, 9, 19-23. Radiochromic film is

widely available, it has very high spatial resolution and has been shown to be suitable

for small field dosimetry provided correcting scanning methodologies are used 10, 19-21,

24, 25, however, requires extensive analysis procedures for accurate measurement.

Optically stimulated luminescence dosimeters (OSLD) are becoming more

widely available in radiation oncology departments for relative dosimetry and in-vivo

dosimetry 26-30. Despite the small size of the OSLDs, 5 mm diameter, there has been

limited investigation of use of OSLDs for small x-ray field dosimetry. In 2013, Pham

et al31 as part of a masters’ thesis investigated making masked OSLDS’s where the

active surface area of the detector was reduced by blocking part with black stickers.

From the thesis, the results were reported that relative output factors could be

adequately determined for cone sizes of 10 mm diameter or greater with the technique

used but not for smaller field sizes. In another study, Yukihara et al. utilized OSLDs

(2 mm diameter by 0.3 mm thickness) to evaluate the relative output factors of Gamma

Knife radiosurgery with 14 mm helmet collimation30. They obtained output factors

with a precision under 1.5% which indicated the potential of OSLDs in small field

dosimetry.

In this work, we developed a novel method to determine relative output factors

for x-ray beam sizes down to 4 mm diameter using the Brainlab SRS cones. We have

constructed and tested a series of variable sized OSLDs with active detector diameters

ranging from 2 mm to 0.6 mm. An experimental extrapolation technique was applied

to minimize the impact of volume averaging and water drops placed into the OSLD

chip in order to remove air gaps around and within the OSL during irradiation.


5.2.1 Construction of modified OSLD

NanoDot (Landauer, Inc., Glenwood, IL, USA) OSLDs were used in this study

to construct the modified dosimeters. An InLight® microStar (Landauer, Inc.,

Glenwood, IL, USA) reader was used to analyse the new OSLD’s.

The nanoDot OSLDs were modified to produce a series of reduced diameter

OSL detectors as shown in figure 1. To perform this procedure, different amounts of

the active crystal were painted black producing varying sized circular active crystal

areas in the center of the OSLD. This was performed by painting around a small drop



of dried PVA glue which was applied to the centre of the OSLD. This was later

carefully removed with a needle leaving the sensitive area exposed and intact. OSLD’s

with active circular diameters ranging from 0.6 mm to 1.8 mm were constructed and

tested. Tests were performed with other masking materials such as black stickers and

paper however, inferior results were obtained and will not be discussed here.

Figure 5-1 A demonstration photo of the modified OSLD’s (middle and right)

comparing with a standard nanoDot OSLD (left).

The readout reproducibility of each chip was tested and compared to standard

NanoDot OSLD values. The reproducibility was found to be within ± 3.1 % (2SD) for

chips between 0.6 mm and 1.0 mm, ± 2.9 % for chips between 1 mm and 2 mm

compared to ± 2.1% for the standard 5mm NanoDot Clinical OSL detectors using our

system.

5.2.2 Small field irradiations

All irradiations were performed with a 6 MV SRS photon beam with a Varian

Novalis Trilogy linear accelerator (Varian Medical Systems, Palo Alto, USA). For

output factor measurements, the modified OSLD’s were placed at 1.5 cm depth, with

source-to-surface (SSD) distance of 98.5 cm as per clinical protocol. Brainlab™ SRS

treatment cones (Brainlab, Munich, Germany), with beam sizes of 4.0, 7.5, 10.0 and

30 mm diameter were tested for output factor using jaw defined field sizes of 5 × 5

cm. The geometry for the measurements matched the criteria specified by Brainlab for

output factor calculations in the iPlan® RT treatment planning software V4.5.4

(Brainlab, Munich, Germany).

The relative output factors were calculated by comparing the measurement value

for a determined cone size to the measurement value for the reference field size. The

reference field size was 10 × 10 cm2, and was measured at depth of 1.5 cm with SSD

of 100 cm. The sensitivity factor for each modified OSLD was determined by

irradiating them to known doses of 100 cGy and 200 cGy using a 10 × 10 cm2 field in



standard calibration conditions. Values for relative output factors were determined by

direct comparison of measure light output for the cone field with the standard field.

Monitor units delivered in each case were chosen to irradiate the modified OSLD’s to

similar absorbed dose levels to the standard to minimize the impact of OSL

supralinearity32. The OSLD’s were read out between 20 minutes and 1 hour after

irradiation to minimize the impact of post irradiation fading27. The relative output

factors as previously determined and reported using Gafchromic EBT3 film (Ashland,

NJ, USA) via an extrapolation technique and the PTW 60019 microDiamond (PTW,

PTW-Freiburg, Germany) dosimeter were used as a reference values for this study33.

Both the EBT3 film and the microDiamond have been well established for small field

dosimetry and have been adopted as reference detectors for such work 19, 34-37.

During the irradiation process, the modified OSLD’s were placed on a dedicated

acrylic tray as shown in figure 2 in between slabs of RMI457 Solid Water with 1.5 cm

of Solid Water was used as build-up and 10 cm for backscatter. The OSLD’s were

fixed in place with tape. The OSLD’s were left opened (which is essential for very

small field irradiations) so that accurate alignment of the active crystal with the beams

central axis was achievable. Furthermore, the acrylic tray slot was filled with water to

remove any air gaps within the phantom and detector and thus loss of backscatter.

Verification was performed to establish that the water did not affect the modified

OSLD’s measurement characteristics compared to dry OSLD’s and results showed

negligible effects when the OSLD’s were kept in water for up to 2 hours as long as

they were dried appropriately before readout.

The diameter of the modified OSLD’s was initially determined with ImageJ

software (National Institutes of Health, USA) via image area analysis. To determine

the estimated output factor, a linear extrapolation technique was used whereby the

OSLD’s measured output factor was plotted against the detectors diameter and



extrapolated to zero diameter. That is, extrapolation to zero volume in order to

minimize the volume averaging effect.

Figure 5-2 A dosimetric the set up for modified OSLD’s with the tray slot filled with liquid

water.

5.3 RESULTS

Figure 5-3 demonstrates the measured relative output factors for a 30 mm

diameter Brainlab cone with the modified OSLD’s of various diameters. Results are

plotted as relative output factor versus the diameter of the modified OSLD active layer.

14 OSLD’s, each with their own unique diameter form the results shown. The results

at each diameter were the average of 5 separate experiments performed for each

modified OSLD. Figure 5-3 also shows a linear extrapolated equation whereby an

estimate of output factor at zero detector size is determined. The extrapolated output

factor was found to be 0.976 ± 0.039 (2SD).

Fig. 5-4, Fig.5-5 and Fig. 5-6 show similar results for the measured relative

output factors for 10 mm, 7.5 mm and 4 mm diameter fields, obtained with the same

modified OSLD’s. These figures also show the estimated extrapolated value of relative

output factor which were 0.871 ± 0.043, 0.828 ± 0.031 and 0.660 ± 0.032 respectively.

Results for measured output factor with varying detector diameter show minimal

differences at larger cone sizes such as 30 mm however show larger differences for the

very small cone sizes like 4 mm. These results are consistent with the significance of

the volume averaging effect in output factor assessment. The extrapolated output

factors estimates were found to be consistent with the values measured with both the

PTW microDiamond detector and EBT3 films within measurement uncertainty for all

cone sizes measured.



Figure 5-3 Measured and extrapolated output factor values for a 6 MV SRS x-ray

beam with the 30 mm diameter Brainlab cone using the modified OSLDs with varying

hole sizes.


beam with 10 mm Brainlab cone using the modified OSLDs with varying hole sizes.

y = -0.0038x + 0.9759

0.9

0.95

1

1.05

1.1

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2

30mm Cone OSL Extrapolation

Ou

tpu

t Fa

cto

r

Modified OSL Diameter (mm)

y = -0.0151x + 0.8709

0.76

0.78

0.8

0.82

0.84

0.86

0.88

0.9

0.92

0.94

0.96

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2


Ou

tpu

t Fa

cto

r





beam with the 7.5 mm Brainlab cone using the modified OSLDs with varying hole

sizes.


beam with the 4 mm Brainlab cone using the modified OSLDs with varying hole sizes.

y = -0.0129x + 0.8275

0.7

0.72

0.74

0.76

0.78

0.8

0.82

0.84

0.86

0.88

0.9

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2

7.5mm Cone OSL Extrapolation

Ou

tpu

t Fa

cto

r


y = -0.0297x + 0.6597

0.55

0.6

0.65

0.7

0.75

0 0.2 0.4 0.6 0.8 1 1.2 1.4 1.6 1.8 2


Ou

tpu

t Fa

cto

r




5.4 DISCUSSION

Very small x-ray fields, such as those found with SRS cones with sizes less than

10 mm, show dose profiles that vary substantially due to lateral charged particle

disequilibrium and can exhibit dose gradients across the whole field3. These effects

can significantly impact on the selection of radiation detector used for the

measurements. A larger detector volume, or in our case, surface area, will influence

the measured output factor to a greater degree for smaller cones with a correlated

detector/field size ratio. This effect is shown in figures 5-3 to 5-6, where the measured

output factors tend to increase towards the reference value as the modified OSLD

diameters decreases i.e. the effect is larger for smaller cones. Using an extrapolation

technique, we were able to remove the influence of volume averaging and estimate the

relative output factor for each cone size from 30 mm down to 4 mm.

In the previous work by Morales et al. it was reported that with an extrapolation

technique, using Gafchromic EBT3 film detector, it was possible to determine relative

output factors for very small SRS field sizes 10. In this study, we found that applying

an extrapolation technique to the modified OSLD measurements, we could obtain

relative output factors that were in agreement with the reference values from both the

PTW 60019 microDiamond and the Gafchromic EBT3 film.

Two essential parameters in our technique for small field dosimetry with

modified OSLD’s was the positioning of the OSLD’s under the SRS beam and the

inclusion of water within the phantom to remove any air gaps during irradiation. As

mentioned, the OSLD’s were irradiated in their open state. This was primarily so that

accurate alignment between the active area of the modified OSLD with the central axis

of the SRS beam. This is especially important for the 4 mm cones. Secondly, when the

OSLD is in its normal closed state, an air gap of approximately 0.2 mm to 0.3 mm

exists above the active crystal to the OSLD casing due to the design of the nanoDot

OSLD. Charles38 et al showed that air gaps of this magnitude can affect dosimetry,

especially for very small fields. The introduction of the water to remove the air gap

provides and essential step in the process. When measurements were performed

without water, the measured outputs were found to be up to 10% lower.

While masking the detector using black colour has been effective in reducing the

region of interest, it can still occur that stray light could potentially generate unwanted

signal and contribute to the noise in the current experimental modification to the OSLD



detector. If the detector could be designed by the manufacturer to resemble the

diameters studied in this work, then one would hope that this source of uncertainty

could be reduced. However, at the moment it is almost impossible to measure

unwanted signals and this forms part of the overall signal noise and uncertainty

included in the values.

This work presents a novel method for use of modified OSLD’s for very small

field output factor dosimetry for SRS. To accurately measure relative output factors

for fields smaller than 10mm, an extrapolation technique and the use of water filling

can be used with the modified OSLD’s. We hope to continue this study and construct

modified OSLD’s with greater positional and higher active crystal size accuracy to

improve uncertainties in measured output factors however preliminary results show

promise for the technique used.

Finally, using the formalism presented in the IAEA TRS 483 Code of Practice,

the correction factors 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 for the modified OSLD’s are close to unity for all field

sizes studied with a maximum difference of 1.5%. Thus, the work presented in this

work shows that further research could be performed on OSLDS with the view to

include OSLDs as part of the IAEA TRS 483 Code of Practice.

5.5 CONCLUSION

A set of modified OSL detectors that have varying surface areas can be used to

measure relative output factors for small field SRS cone beams. An extrapolation

technique to zero volume along with water filling around the detector to remove air

gaps are essential components for accuracy for fields smaller than 10 mm diameter.

Results showed that, using this technique that relative output factors measurements

within ±2 % of corrected microdiamond and EBT3 film results were achievable. As

such, standard Nanodot OSLD’s could be modified and used to verify very small field

output factors used in SRS radiotherapy.


No conflict of interest



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93




Chapter 6: A comparison of surface doses for very small field x-ray beams: Monte Carlo calculations and

radiochromic film 95

Chapter 6: A comparison of surface doses

for very small field x-ray beams:

Monte Carlo calculations and

radiochromic film

Overview

Surface dose or skin dose is important factor in radiation therapy. There is a

correlation between skin dose and toxicity during early stage effects such as erythema.

There can be occasional late effects such as Telangiectasia as well. However, it is well

known that treatment planning system in radiotherapy and radiosurgery do not predict

the dose near the surface or at shallow depths, and sometimes in vivo measurements

on the patient are required. There is additional uncertainty regarding skin dose for

small radiation beams, as there is limited published data, and most of the published

data is based on MLC shaped fields. In this study, the skin dose for small circular cones

was calculated through Monte Carlo methods and compared against radiochromic film

measurements. This paper is one of the first to provide skin dose values for cone-based

stereotactic radiotherapy at the ICRP depth of 70 m

96 Chapter 6: A comparison of surface doses for very small field x-ray beams: Monte Carlo calculations and

radiochromic film


Title: A comparison of surface doses for very small field x-ray beams: Monte

Carlo calculations and radiochromic film measurements

Authors: Johnny E Morales, Robin Hill, Scott B. Crowe, Tanya Kairn, J.V.

Trapp




Scott B. Crowe


Robin Hill


Tanya Kairn


J.V. Trapp




Status: Published 20 March 2014


SCOPUS Authors h-index: 6

https://doi.org/10.1007/s13246-014-0260-2



ABSTRACT

Stereotactic radiosurgery treatments involve the delivery of very high doses for

a small number of fractions. To date, there is limited data in terms of the skin dose for

the very small field sizes used in these treatments. In this work, we determine relative

surface doses for small size circular collimators as used in stereotactic radiosurgery

treatments. Monte Carlo calculations were performed using the BEAMnrc code with

a model of the Novalis Trilogy linear accelerator and the BrainLab circular

collimators. The surface doses were calculated at the ICRP skin dose depth of 70 m

all using the 6 MV SRS x-ray beam. The calculated surface doses varied between 15

and 12% with decreasing values as the field size increased from 4 to 30 mm. In

comparison, surface doses were measured using Gafchromic EBT3 film positioned at

the surface of a Virtual Water phantom. The absolute agreement between calculated

and measured surface doses was better than 2.0% which is well within the uncertainties

of the Monte Carlo calculations and the film measurements. Based on these results, we

have shown that the Gafchromic EBT3 film is suitable for surface dose estimates in

very small size fields as used in SRS.


radiochromic film

6.1 INTRODUCTION

Stereotactic radiosurgery (SRS) involves the delivery of a high radiation dose,

typically 15 to 20 Gy, using small size radiation beams for treatments of lesions within

the brain 1. These lesions are usually malignant brain metastases or benign

arteriovenous malformations (AVMs) and require high spatial and dosimetric accuracy

due to the need for accurate delivery to the small lesion as well as minimising radiation

dose to other tissues such as the brain stem, optic chiasm and other critical structures.

SRS treatments usually involves a number of beams using arcs and/or conformal fields

in order to achieve an optimum dose distribution and normal tissue doses.

Skin toxicity can also be a dose-limiting factor for radiotherapy treatment

planning, despite the skin-sparing effect of megavoltage x-ray dose build-up2-4. Skin

dose is a particular concern in SRS treatments, due to the high single-fraction doses

delivered, and there is a growing interest in the incidence of skin toxicity associated

with stereotactic body radiotherapy (SBRT or SABR)5.

The ICRP defines the skin depth to be 70 m which corresponds to the basal cell

layer thickness6. Surface doses for megavoltage x-ray beams have been measured

using a number of different detectors including parallel-plate ionisation chambers,

thermoluminescent dosimeters (TLDs), metal-organic semiconductor field-effect

transistors (MOSFETs) and various types of radiochromic film2, 3, 7-20. While some

parallel-plate ionisation chambers, such as the Attix chamber are well characterised

for dose measurements in the build-up region, they are not suitable for in-vivo

dosimetry measurements2, 3, 21. In addition to measurements, surface doses for x-ray

beams can also be determined by Monte Carlo calculations. Monte Carlo methods are

regarded as the gold standard for accurate dose calculations of ionising radiation

beams22-24 and there have been a number of studies which have used Monte Carlo

methods for determined surface doses for conventional field sizes2, 4, 25-28.

To date, there is limited data for surface doses for the very small field sizes as

typically used in the delivery of SRS treatments4, 29. Paskalev et al examined the

dosimetry of SRS beams, including surface doses measured with radiographic film,

generated from a 10MV x-ray beam and using 1.5 and 3 mm circular collimators30 and

found that the surface doses decreased slightly with increasing size of the collimator.

In a comprehensive study of the dosimetry of the Brainlab m3 micro-MLC system,



Ding et al used both experimental and Monte Carlo methods to characterise the

dosimetry29. As part of their results, the Monte Carlo depth dose data included surface

doses for field sizes ranging from 6×6 to 100×100 mm2 with relative surface dose

values ranging between 26 – 32% of Dmax.

The purpose of this study is to determine surface doses for a variety of very small

size fields as defined by circular applicators ranging from 4 to 30 mm as used in SRS

treatments. Surface doses are calculated using Monte Carlo modelling of the linear

accelerator using the BEAMnrc Monte Carlo code. In addition, surface doses were

measured for these applicators using the Gafchromic EBT3 radiochromic film.


In this work, surface doses were determined for the 6 MV SRS x-ray beam as

produced by a Novalis Trilogy linear accelerator (Varian Medical Systems, Palo Alto,

USA) which has a thin flattening filter in order to produce a dose rate of 1000 MU per

minute31-33. Beam collimation for the SRS x-ray beams was achieved by using the

BrainLab circular applicators (BrainLab, Germany) with diameters of 4, 7.5, 10, 20

and 30 mm diameter as defined at the isocentre. The X and Y collimator jaws were set

to 5 cm for all measurements and calculations. These jaw sizes were used during

commissioning of this linear accelerator but it is noted that a recent notification from

the manufacturer recommends different jaw settings for each circular collimator

(Safety Notice 09-06-29.BAV.2).

6.2.1 Commissioning Monte Carlo model

The 6MV SRS x-ray beam was modeled using the BEAMnrc/EGSnrc Monte

Carlo code (Version 4, release 2.3.2)23, 34, 35. Simulations were performed on

supercomputing facilities at the Queensland University of Technology. The

supercomputer at this location has 1924 64-bit Intel Xeon Cores. All the specifications

for the geometry and materials used within the linear accelerator model including the

special SRS flattening filter were supplied by Varian Medical Systems. The adjustable

incident electron beam parameters were optimized to an elliptical beam with Gaussian

distributions of X = 0.4 cm and Y = 0.4 cm in FWHM and with no angular spread

striking the tungsten target along the central axis. The optimization process involved

the tuning of the initial electron energy to match the PDD curve for a 10 × 10 cm2 open

field size. A reduced Chi squared method was applied to every PDD obtained by


radiochromic film

Monte Carlo. Energies were varied between 5.0 and 7.0 MeV in steps 0f 0.2 MeV. The

reduced Chi squared results were plotted versus energy and a minimum was reached.

This minimum value was taken to be the best value for Monte Carlo simulations. In

our case the final value for the incident electron energy was 6.06 MeV.

For both the BEAMnrc and DOSXYZnrc user codes, the electron cut off energy

(ECUT) and the photon cut off energy (PCUT) were set to 0.521 MeV and 0.010 MeV

respectively. PRESTA-I and PRESTA-II were turned on for the boundary crossing

algorithm (BCA) and electron-step algorithm respectively. The directional

bremsstrahlung splitting (DBS) variance reduction technique was also used with

splitting factor of 100036, 37.

The initial testing of the BEAMnrc model was for an open field size of 10 × 10

cm2 and no circular applicator. A total of 4×107 histories were used. The initial model

of the x-ray beam was verified by comparison of measured and calculated percentage

depth dose data and cross-plane beam profiles as measured in a PTW MP3 water tank

(PTW, Freiburg, Germany). The procedure for comparing data involved working out

the absolute difference by subtracting the measured data from the Monte Carlo data at

each measurement point. For the 10 × 10 cm2 open field size, depth doses and cross

plane profiles were measured with a PTW Advanced Markus parallel-plate ionisation

chamber (PTW, Freiburg, Germany) and PTW 60003 diamond detector (PTW,

Freiburg, Germany).

6.2.2 Surface Dose simulations for circular collimators

After verifying the Monte Carlo model for the 10 × 10 cm2 open field size, phase

space files were generated for the circular collimators with diameters of 4, 7.5, 10, 20

and 30 mm using the BEAMnrc code. Percentage depth doses were then calculated for

each circular collimator using the DOSXYZnrc user-code (V4 r2-3-0) and compared

with PDDs measured with a PTW 60012 diode (PTW, Freiburg, Germany) at a source

surface to distance of 100 cm. Voxel sizes of 0.25×0.25×0.1 cm3 were used to score

the dose for central-axis PDDs. A total of 8×109 incident particles were used to obtain

statistical uncertainties of less than 2 %. For consistency, the same user-specified

simulation parameters were used as in the phsp file calculation. To model electron

transport as accurately as possible, a global ECUT of 0.521 MeV was specified which

corresponds to an electron range of approximately 2.5 m in water (NIST



DATABASE at http://physics.nist.gov/cgi-bin/Star/e_table.pl). The EXACT

boundary crossing algorithm was turned on for surface dose calculations38.

For the surface dose calculations with the circular collimators, additional

calculations were carried out using the DOSXYZnrc user-code. The dose was scored

in voxels with thicknesses of 10 m for the first 0.1 mm depth, which includes the

clinically relevant skin depth of 70 m6, as well as at depths of 1 mm, 2 mm and dmax

= 1.5 cm. All surface doses scored at or close to the surface of the phantom were

normalized to Dmax.

6.2.3 Surface Dose measurements using Gafchromic EBT3 film

All measurements were performed with Gafchromic EBT3 radiochromic film

(Ashland, Wayne, USA) which has a number of advantages over the earlier versions

of this film including a polyester substrate which prevents the formation of Newton’s

rings and the use of a symmetrical structure for the different layers in the

manufacturing of the film 39, 40. The EBT3 film has the active layer of 30 µm thickness

which is located between the two polyester layers of 125 µm thickness. The EBT3 film

sheet was cut into 2 × 2 cm² pieces for both the dose calibration and surface dose

measurements. The film piece was positioned at the surface and at a depth of Dmax

between the Virtual Water blocks41. For each set of measurements, a minimum of two

EBT3 film pieces were used in this study, and the mean dose absorbed by these film

pieces was used for analysis purposes.

The process for preparing, reading out the films and analyzing the dose

information was consistent with manufacturer’s recommendations. All films were read

out on an EPSON 10000 XL (EPSON) flatbed scanner maintaining a fixed orientation

of the film during the readout. The orientation chosen was landscape which keeps the

short edge of the film parallel to the scan direction. The EPSON scan software package

was used to scan films using the red channel only at a resolution of 150 DPI in

transmission mode with all image adjustment features switched off. The dosimetric

analysis was performed using the RIT software package V5.2 (Radiological Imaging

Technology, Inc, USA).

The RIT software needs a calibration curve which allows the conversion of the

pixel value to absorbed dose. This was achieved by irradiating the film pieces with

known doses of 0 – 3 Gy by using a 10 × 10 cm2 field size and the films were located

http://physics.nist.gov/cgi-bin/Star/e_table.pl


radiochromic film

at the depth of dmax. A median filter of 3 × 3 pixels was applied to all the scanned

images. All films were read out approximately 24 hours after their irradiation.

For the surface dose measurements with the different collimators, the EBT3 film

piece was placed at the surface of the Virtual Water phantom at an SSD of 100 cm.

The uncertainty in the EBT3 film surface dose measurements was determined to

be 2 % (1 Standard Deviation) using the ISO GUM methodology42-44. The factors

contributing to uncertainties in surface dose measurements included: variation within

the OD measurements for the pixels in the region of interest, variations due to film

non-uniformity as well as uncertainties in the curve fit for the EBT3 film calibration

curve.

6.3 RESULTS AND DISCUSSION

6.3.1 Commissioning Monte Carlo model

The uncertainty in the Monte Carlo calculated depth doses was less than 1% as

determined within the DOSXYZnrc user code. Figures 6-1 and 6-2 show the

comparison between the Monte Carlo calculated and measured central axis depth dose

curves and cross-plane profiles respectively for the open 10 × 10 cm2 reference field.

The agreement between the Monte Carlo calculated and the measured doses for the

PDD shown in figure 1 was within 1% for depths of 1.0 to 30.0 cm and within 2% for

depths within 0.5 to 0.99 cm, which is well within an acceptance criteria of 2%45, 46.

Measurements with the Advanced Markus for depths between 0 and 0.5 cm are not

shown as they were deemed unreliable due to the meniscus effect near water surface.



Figure 6-1 Percentage depth dose in water calculated for a 10 × 10 cm2 field.

BEAMnrc/DOSXYZnrc versus measurement by an Advanced Markus ionisation

chamber.

Figure 6-2a shows the Monte Carlo calculated and measured profiles at three

depths in water across the X jaws. The depths were 1.4, 10 and 20 cm. All profiles

were normalized with their respective PDD value at each depth. Figure 6-2b shows the

difference between calculated and measured data. A maximum absolute difference of

1.4% was obtained at the penumbra regions.

The dose distance to agreement (DTA) used in the evaluation was 2%/1mm

which is very consistent with tolerances for very small fields.

Please note that although the model was fine-tuned at 10 × 10 cm2, however, the

PDD measurements for the 4, 10, 20 and 30 mm cones presented in Figure 6-3 show

that the modelling performed with Diode does in fact agree well with the small field

PDDs with an uncertainty of 2 %. The uncertainty refers to Type-A uncertainty for

both measurements and Monte Carlo calculation.


radiochromic film

Figure 6-2 a) Cross profiles at depth in water calculated by BEAMnrc/DOSXYZnrc

for a 10 × 10 cm2 field versus measurement with a diamond at depths of 1.4, 10 and

20 cm. b) Absolute difference between calculation and measurement for each depth.

6.3.2 Surface Dose simulations

The comparison between the Monte Carlo calculated and measured central axis

depth dose curves for the circular collimators are shown in Figure 6-3. For all the

circular collimators, the agreement between Monte Carlo calculated and measured

percentage depth doses was within 1% for depths between 0.5 and 30 cm. This is well

within an acceptance criteria of 2 %45, 46.



Figure 6-3 Percentage depth doses calculated by BEAMnrc/DOSXYZnrc and

measured with a diode for: a) 4 mm circular collimator, b) 10 mm circular collimator,

c) 20 mm circular collimator and d) 30 mm circular collimator.

6.3.3 Surface Dose measurements using Gafchromic EBT3 film

A summary of the relative surface doses for the Monte Carlo calculations and

the Gafchromic EBT3 measurements is presented in Table 6-1. The estimated

uncertainty in the Gafchromic EBT3 film measurements was 2.0%.


radiochromic film

Table 6-1 Relative surface doses for Brainlab SRS circular collimators determined

by Monte Carlo calculations and Gafchromic EBT3 measurements

Circular collimator

diameter (mm)

Monte Carlo

surface dose

(% of Dmax)

EBT3 film

surface dose

(% of Dmax)

4 15.0 14.5

7.5 12.8 15.5

10 12.3 15.0

20 12.3 13.0

30 11.9 14.0

For the Monte Carlo calculations, the surface dose calculations range from

15.0% for the 4 mm applicator to 11.9% for the 30 mm applicator. It is of interest to

note that there is a decrease in the surface dose by just over 3% from the smallest to

the largest cone. However, this result is consistent with previous studies that reported

decreases in the relative surface dose while using small field sizes and for increasing

field sizes29, 30. The study by Ding et al found that the Monte Carlo calculated surface

doses varied from 28% for the 6×6 mm2 field, 26% for the 12×12, 18×18 and 24×24

mm2 fields and then started increasing from a field size of 30×30 mm2 29. Similarly,

the study by Paskalev et al found a reduction in the relative surface dose from the 1.5

and 3 mm diameter fields31. While it is accepted that surface dose increases as a

function of field size for megavoltage x-ray beams due to increasing scatter, most

studies do not use very small field sizes of less than 3 cm for which lateral dose

equilibrium is achieved4.

The measured surface doses for the five circular collimators had relative surface

values in the range between 13.0 – 15.5%. These measured values were in agreement

with the Monte Carlo calculated doses to within 3% measurement uncertainty. These

results show that the EBT3 film can be used for surface dose estimation in SRS beams

and would have application in either benchmark relative dosimetry measurements or

for in-vivo surface dose measurements.

It is important to know the skin dose as accurately as possible because skin

toxicity can be a problem in SBRT treatments5. It is important to note that there have

been some differences reported between MOSFETs and radiochromic film for breast

radiotherapy treatments47. Furthermore, treatment planning systems may not

adequately deal with skin dose calculation making it harder to know the exact skin

dose for some treatments like early-stage non-small-cell lung cancer5. Therefore,



having an independent method to verify the skin dose is definitely important in order

to achieve optimal skin sparing. Our model is one step forward in having such tool to

evaluate skin toxicity in treatments where conical collimators are used in SRS and

SBRT treatments.

6.4 CONCLUSIONS

This study has determined the surface doses for a 6MV SRS x-ray beam for very

small field sizes from circular collimators with diameters ranging from 4 – 30 mm.

The Monte Carlo method showed that the surface doses were: 15.0% for 4mm, 12.8%

for 7.5 mm, 12.3% for 10 mm, 12.3% for 20 mm and 11.9% for 30 mm. The EBT3

film measurements produced the following surface doses: 14.5% for 4 mm, 15.5% for

7.5 mm, 15.0% for 10 mm, 13.0% for 20 mm and 14.0% for 30 mm. The uncertainty

for EBT3 film measurements was 2% (1 Standard Deviation). This work has shown

that both methods gave consistent surface dose values and indicates that Gafchromic

EBT3 film can be used for surface dose measurements in radiotherapy departments

where Monte Carlo simulations are not available for stereotactic radiosurgery beams.


Computational resources and services used in this work were provided by the

High Performance Computing and Research Support unit, Queensland University of

Technology, Brisbane, Australia. Also, we would like to acknowledge that Dr S. B.

Crowe’s contribution to this work was supported by the Australian Research Council

through Linkage Grant No. LP110100401.



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111




Chapter 7: A study of dose inhomogeneity correction in a commercial treatment planning system for stereotactic

radiosurgery 113

Chapter 7: A study of dose inhomogeneity

correction in a commercial

treatment planning system for


Overview

Accurate treatment delivery is reliant on accurate dose calculations in the patient.

For cranial stereotactic radiotherapy, treatment fields may traverse through significant

volumes of low or high density media, including regions of air cavities. The study of

these inhomogeneity effects in radiation dosimetry is not new. However, most work

has been concentrated on using multi-leaf collimator (MLC) shaped fields. This study

investigated the accuracy of the Clarkson pencil beam algorithm in the iPlan TPS for

calculating dose in low and high density media using the Brainlab circular cones of 4,

7.5 and 10 mm diameter, by comparison against full Monte Carlo calculations. This

comparison was performed in slab phantoms containing tissue inhomogeneities as well

as in an anthropomorphic head and neck phantom. This work also shows the use of a

freeware package from McGill, Canada, namely, the McGill Monte Carlo Treatment

Planning (MMCTP) platform. MMCTP, according to its authors, is able to facilitate a

systematic, platform-independent, large-scale MC treatment planning for different

treatment sites by creating BEAMnrc/DOSXYZnrc input files which are then used for

large scale parallel processing calculations on large computer clusters or High

Performing Computing servers. This paper is a pilot study on the implementation of a

Monte Carlo based independent check system for iPlan patient plans in the clinic.

114 Chapter 7: A study of dose inhomogeneity correction in a commercial treatment planning system for



Title: A study of dose inhomogeneity correction in a commercial treatment

planning system for stereotactic radiosurgery

Authors: Johnny E Morales, Martin Butson, Robin Hill, Scott B. Crowe, Jamie

V. Trapp




Martin Butson


Robin Hill


Scott B. Crowe


Jamie V. Trapp





radiosurgery 115

Abstract

Small x-ray fields are used in complex radiation treatment fields for stereotactic

radiosurgery (SRS) by arc or IMRT delivery. Accurate treatment delivery is reliant on

accurate dose calculations to the region of interest containing the tumour volumes

usually within the brain. In some cases, the treatment fields may traverse through

significant amount of low or high density tissue including regions of air cavities. The

purpose of this work was to investigate the accuracy of the Brainlab iPlan planning

system with the Clarkson pencil beam algorithm in planning with Brainlab circular

cones with diameters of 4.0, 7.5 and 10.0 mm. Doses were calculated using iPlan in

slab phantoms containing low and high density tissue equivalent materials as well as

in a head anthropomorphic phantom. For comparison, doses were also calculated for

the same geometries using the BEAMnrc/DOSXYZnrc Monte Carlo codes using

phase spaces for the circular cones all within the MMCTP Monte Carlo framework. In

the low density slab phantoms, the Monte Carlo calculated doses had percentage

differences in predicted dose of 48%, 39% and 32% for the 4 mm, 7.5 mm and 10 mm

diameter cones respectively. In comparison, the doses in the high density tissue was

up to 3.5% higher as compared to the doses calculated by iPlan. For the DVH

calculations to a 4 mm target volume within the middle of the head phantom, the value

of V100 was more than 12% less than the value calculated within iPlan. Based on these

results, care needs to be taken when planning SRS treatments with large tissue

inhomogeneities particularly in very low density tissues. In addition, the expected dose

differences in tissue inhomogeneities can be estimated form the Monte Carlo results

in this study.



7.1 INTRODUCTION

Radiotherapy involves the delivery of high doses of radiation to the tumour

volume while minimising the dose to healthy tissues. An important component of

radiotherapy is the individualised patient treatment plan calculated by a treatment

planning system (TPS). Current commercial treatment planning systems include

Monte Carlo1-4 calculation engines which are designed to calculate dose for complex

treatment fields shaped with multi-leaf collimators (MLCs) as used for Intensity

Modulated Radiotherapy (IMRT), dynamic conformal arcs and Volumetric Modulated

Arc Therapy (VMAT) including treatments involving stereotactic radiosurgery (SRS)

and stereotactic body radiation therapy (SBRT).

One of the commercial treatment planning systems for SRS dose calculations is

the Brainlab iPlan system (Brainlab, AG, Feldkirchen, Germany) which has two dose

calculation algorithms available: a Monte Carlo (MC) algorithm used for planning

MLC shaped beams, and a Clarkson pencil beam algorithm used for planning with the

Brainlab circular cones. The MC algorithm is based on the X-ray Voxel Monte Carlo

dose algorithm developed by Fippel and Kwarakow5-7. The Monte Carlo algorithm is

only available for planning with MLC shaped beams and currently does not support

any planning calculations with Brainlab circular cones.

There have been a number of studies which have examined the accuracy of the

TPS dose calculation algorithms in calculating the dose to the different tissues in the

body including low density structures8-13. These studies have used a combination of

block phantoms containing slabs of lung type material and/or in anthropomorphic

phantoms with lung equivalent inserts.

Although the CyberKnife system has two dose calculation algorithms, a Ray

Tracing algorithm and a Monte Carlo algorithm, Liang et al14 developed their own

treatment planning algorithm for their CyberKnife system to improve the lateral

scatter in irregularly shaped small field for the CyberKnife system. They achieved this

by fitting the kernel and intensity profile to the commissioning data. A validation

against the Ray Tracing and the Monte Carlo algorithm was performed and they found

that their algorithm improved the accuracy for treatment plans on the CyberKnife

system.


radiosurgery 117

Oyoke et al15 performed a study to determine whether CyberKnife (CK)-based

SBRT using a Ray Tracing (RyTc) algorithm is comparable dosimetrically to that

of Monte Carlo (MC) for thoracic spinal lesions. Their study included 37 patients and

the comparisons were made using RyTc and MC. They found that the RyTc algorithm

overestimates the MC calculated average percentage volume of PTV covered by the

prescribed dose and have unpredictable effects on doses to organs at risk, particularly

the spinal cord. They recommended that the use of the RyTc algorithm should be

limited and should always be verified by the MC algorithm.

Song et al16 performed a study on comparing clinical outcomes in SBRT for lung

tumours between Ray-Tracing and Monte Carlo algorithms. Their study included 35

patients who received SBRT treatment on a CyberKnife machine. It was found that the

response rate for Ray-Tracing algorithm was 77.3 % compared to the Monte Carlo

algorithm with 100 %. They concluded that the clinical outcome and toxicity of lung

SBRT between the Ray-Tracing and Monte Carlo algorithms were similar except for

the response rate when the same apparent doses were prescribed. They concluded that

the lower response rate in the Ray-Tracing group, however, did not compromise the

local control rates. Their recommendation was that reducing the prescription dose for

Monte Carlo algorithm may be performed but done it should be done with caution.

This was a significant conclusion as there have been questions about modifying the

dose produced by Ray-Tracing versus the dose produced by Monte Carlo algorithms

in general.

Another study by Pan et al17 involved testing the accuracy of the Monte Carlo

algorithm in the MultiPlan planning system for CyberKnife. In this system, a plan is

first calculated using a simple Ray Tracing algorithm and then re-calculated using the

Monte Carlo algorithm. They produced single beam plans on a solid water phantom

and then created plans on a thorax phantom to mimic a clinical patient treatment plan.

Their results found that for lung cases, the gamma passing rate for Monte Carlo was

98.31 % and 97.28 % for homogeneous and heterogeneous geometries. In contrast, the

Ray Tracing algorithm had a passing rate of 79.25 % for the heterogenous situation.

Thus, they recommended to use the Monte Carlo algorithm for cases involving

heterogenous media.

For smaller x-ray fields as typically used in SRS treatments, there have been less

investigations of the low density tissue structures18-20, for example Jones et al20



performed a study of dose inhomogeneity correction algorithms for small fields down

to a minimum field size of 5 mm. They argued that historically the field sizes used in

radiotherapy have been larger than 5 × 5 cm2 but with the introduction of IMRT and

other modulated techniques, there is a greater need to investigate fields smaller than 5

× 5 cm2. They performed Monte Carlo calculations using the BEAMnrc/DOSXYZnrc

package for small fields and these Monte Carlo calculations were used as the

benchmark in their study20.

Clinical areas of interest for SRS treatments where low and high density tissues

occur include sites such as the nasopharyngeal area and the skull. Small tumours

located in this vicinity may require treatment through or near low density regions. As

a consequence, the planning algorithm used can have an impact on calculated dose and

dose coverage.

In this study, we investigate the Clarkson pencil beam algorithm in iPlan for

calculating dose in low and high density media using the Brainlab circular cones of 4,

7.5 and 10 mm diameter in comparison with full Monte Carlo calculations using

previously validated phase space files for the Brainlab circular cones21. The dose

comparison was performed in slab phantoms containing tissue inhomogeneities as well

as in an anthropomorphic head and neck phantom.


7.2.1 Plans calculated in iPlan RT Dose

Plans were calculated on iPlan RT Dose version 4.5.3 using an in-house Virtual

Water phantom containing slabs of low and high density materials. This phantom is

shown in Figure 7-1 below. This phantom provided a Virtual Water depth of 7 cm

before a 3 cm region of low or high density inhomogeneity. The low density region

was produced by a CIRS slab (CIRS, Norfolk, VA, USA) which had a mass density of

0.3 g cm3. The high density region was produced by a CIRS slab (CIRS, Norfolk, VA,

USA) which had a mass density of 2.6 g cm3. The slabs had square dimensions of 30

× 30 cm2. These were taken as representative of low and high density tissue but it is

acknowledged that the human body has a wide range of densities. The iPlan RT Dose

planning software used a Hounsfield Unit to electron density curve conversion curve

for dose calculation. The CT scans were acquired on a Toshiba LB Aquilion CT

scanner (Toshiba, Japan) using a CT slice thickness of 0.5 mm.


radiosurgery 119

Figure 7-1 In-house Virtual Water phantom showing a low density slab inserted in the

middle section

The linear accelerator used for this study was a Novalis Trilogy (Varian Medical

Systems, Palo Alto, USA) using a 6 MV SRS beam and equipped with Brainlab

circular cones (Brainlab, AG, Feldkirchen, Germany) of 4, 7.5 and 10 mm diameter.

This linac can deliver radiation at a dose rate of 1000 MU/min and is configured in the

iPlan planning system using the Clarkson pencil beam algorithm for dose calculation

with the cones. This algorithm calculates dose based on tissue phantom ratios, relative

output factors and single beam profiles measured in water with no correction for

inhomogeneous tissue. The data used in this algorithm was measured according to the

manufacturers specifications. The specifications relevant to this thesis were as follows:

The Source to Surface Distance (SSD) for measurements of Percentage Depth Doses

(PDDs) and Output Factors was 98.5 cm, the depth of measurement for output factors

was 1.5 cm. The output factor measurements and PDDs were performed using a PTW

60012 Diode E as well as the Off Axis Ratios. These guidelines were followed to

comply with the manufacturer recommendations although the knowledge in this thesis

would recommend to use more modern techniques.

PDDs and cross profiles at depth were calculated for the three Brainlab circular

cones. The beams were incident at a gantry angle of 0 degrees, a collimator angle of

90 degrees and isocentre depth of 1.5 cm with a dose of 2.0 Gy prescribed at the

isocentre. The source to axis distance (SAD) was 100 cm. The PDDs were calculated



from the top of the phantom to the bottom and the cross profiles at depth were

calculated laterally through the middle of the low or high density slab.

7.2.2 Plans calculated with Monte Carlo in MMCTP

The Monte Carlo software package used in this work is the McGill Monte Carlo

Treatment Planning (MMCTP) package22. The MMCTP software is an software

platform which allows the import and export of DICOM files for the purpose of

treatment planning using full Monte Carlo dose calculations with the BEAMnrc Monte

Carlo system23. MMCTP is able to generate BEAMnrc/DOSXYZnrc23, 24 input files

from the DICOM format files which are necessary for calculating Monte Carlo dose

distributions using BEAMnrc/DOSXYZnrc. MMCTP is also capable of comparing

multiple dose distributions from different planning systems provided they are in

DICOM RT format. This platform simplified the workflow for our study as shown in

Figure 7-2.


radiosurgery 121

Figure 7-2 Workflow diagram showing the process followed for comparing iPlan

generated treatment plans with Monte Carlo generated plans using MMCTP platform.

To calculate Monte Carlo plans, shown in figure 1, the plans created in iPlan RT

Dose 4.5.3 (Section 2.1) were exported in DICOM format to MMCTP. The DICOM

export from iPlan included the CT images, target structure and the plan parameters

including the cone size, Source to Surface Distance (SSD), depth, gantry angle,

collimator angle, field size and Monitor Units. MMCTP converted the DICOM data

into an MMCTP specific format for subsequent processing. MMCTP created input



files that were compatible for Monte Carlo calculations using BEAMnrc and

DOSXYZnrc programs which are part of the EGSnrc Monte Carlo system. The

BEAMnrc code produced a phase-space file simulating the head geometry and beam

characteristics which was then used by DOSXYZnrc for calculating dose in the in-

house phantom. The head geometry and Brainlab circular cone parameters for the

Novalis Tx linear accelerator were similar to those published earlier21.

For calculations in the in-house Virtual Water phantom, MMCTP created a

phantom file in egsphant format necessary for dose calculation in DOSXYZnrc. This

file was created from a specific PEGS4 data file which accounted for Hounsfield Units

and the atomic composition of different media. Comparison of the dose distributions

produced by iPlan planning systems was performed by comparing the percentage

depth dose curves and cross profiles for the 4, 7.5 and 10 mm diameter Brainlab cones.

7.2.3 Clinical treatment plans calculated on an anthropomorphic phantom

Three clinical SRS plans were calculated using iPlan on a RANDO head and

neck anthropomorphic phantom as shown in Figure 7-3. Each plan was prepared with

a single circular cone. A representative tumour volume, being the PTV used for

prescribing the dose, was drawn inside a water equivalent material region surrounded

by a region of low density tissue. The representative tumour volumes being 0.372 cm3,

0.081 cm3, 0.012 cm3 for the 10 mm, 7.5 mm and 4 mm diameter cones respectively.

The tumour shape was a sphere in shape. A dose of 2.0 Gy per fraction was prescribed

to the 80 % isodose line. Each plan had 5 arcs with a gantry angle range of 160 degrees

as per our clinical program and the couch angle range was 135 degrees also as per our

clinical program. An assessment of the amount of low density material which each arc

was traversing was performed by obtaining the path length versus the water equivalent

path length. It was found that on average the path length was 1.4 cm larger than the

water equivalent length. This would indicate that relative amount of low density which

each of the arcs was passing through in each of the plans. A similar process was

repeated for the 7.5 and 10.0 mm diameter cone. The CT slice thickness was 0.5 mm

to maximise the spatial resolution for all cone diameters. Figure 3 shows the dose

distribution on three CT slice planes of the anthropomorphic phantom as displayed in

MMCTP. Once the iPlan plans were calculated, the DICOM RT files including the CT

images, structures, MU, prescription dose were exported into MMCTP. Dose Volume


radiosurgery 123

Histograms (DVH) were also calculated within the MMCTP software for each clinical

plan.

Figure 7-3 Anthropomorphic phantom showing treatment plan with 7.5mm Brainlab

cone

7.3 RESULTS

7.3.1 Depth doses for low and high density material slab

Figures 7-4 and 7-5 shows the percentage depth dose calculated by iPlan RT

Dose and by DOSXYZnrc for the low and high density slabs. In Figure 7-4, the iPlan

RT Dose calculated PDD and the Monte Carlo calculated PDD both agree within 0.5%

for the region from zero depth to the slab inhomogeneity. There was a reduction in

percentage depth dose value for each of the circular cones at the region of slab

inhomogeneity. The reduction in percentage dose value was 28%, 22% and 20% for

the 4, 7.5 and 10 mm respectively. These values relate to percentage differences in

predicted dose of 48%, 39% and 32% for the 4 mm, 7.5 mm and 10 mm cone

respectively. It can be seen that the Clarkson based algorithm for the circular cones in

iPlan RT Dose shows little change. The general trend for the Monte Carlo calculated

PDD is such that as the field size is decreased the reduction is larger. It can also be

seen that in the region beyond the low density material electronic equilibrium is re-

established and the agreement between iPlan and Monte Carlo calculations is within

0.5% once again. In Figure 7-5, the iPlan RT Dose calculated PDD and the Monte

Carlo calculated PDD both agree within 0.5% for the region from zero depth to the

high density inhomogenenity. In the region of high density, the PDD value seems to



increase slightly as the field size is reduced to 4 mm diameter. The corresponding

increase values were 3.5%, 1.3% and 1.1% for the 4, 7.5 and 10 mm respectively.

Figure 7-4 Percentage depth dose calculated by iPlan and by Monte Carlo for the

Brainlab circular cones in a virtual water phantom containing a low density region.


radiosurgery 125

Figure 7-5 Percentage depth dose calculated by iPlan and by Monte Carlo for the

Brainlab circular cones in a virtual water phantom containing a high density region.

7.3.2 Profiles for low and high density material

Figures 7-6 and 7-7 shows the cross profiles calculated in iPlan RT Dose and

calculated by DOSXYZnrc at geometrical centre of the low and high density slab

material, respectively. These profiles were imported into MMCTP for display and

comparison. The profiles are normalised to 100 % for the value on the central axis at

this depth. It can be appreciated that the penumbra of the profile calculated by Monte

Carlo is broader than the profile calculated in iPlan for the same cone and is more

pronounced at the edge of the beam. This would have an impact on Gradient Index for

each plan. The Gradient Index is defined as the ratio of the volume of half the

prescription isodose to the volume of the prescription isodose25.



Figure 7-6 Cross profiles at depth calculated by iPlan and by Monte Carlo for the

Brainlab circular cones in a slab of low density material inserted in virtual water

phantom

Figure 7-7 Cross profiles at depth calculated by iPlan and by Monte Carlo for the

Brainlab circular cones in a slab of high density material inserted in virtual water

phantom


radiosurgery 127

7.3.3 Full width at half maximum distances for the low and high density

Table 1 below shows the full width at half maximum (FWHM) distances for the

low and high density slabs. The FWHM is defined at values of 90 – 10%, FWHM90-

10, as well as 80 – 20% or FWHM80-20. The table shows that the low density slab

broadens the penumbra by approximately 1 mm for the 4 mm cone, 2 mm for the 7.5

mm cone and 3 mm for the 10 mm cone respectively.

Table 7-1 FWHM90-10 and FWHM80-20 values for the High and Low density slab materials

for the 4, 7.5 and 10 mm cones.

Monte Carlo – Low density iPlan – Low density

Cone (mm) FWHM90-10 FWHM80-20 Cone (mm) FWHM90-10 FWHM80-20

4 3.7 1.7 4 2.6 1.6

7.5 5.3 2.5 7.5 3.1 1.9

10 6.4 3.1 10 3.4 2.0

7.3.4 DVH comparison for Anthropomorphic plans

Figure 7-8 shows the Dose Volume Histogram (DVHs) for the plans calculated

in the anthropomorphic head phantom using the 4, 7.5 and 10 mm diameter cones. The

curves plotted correspond to the representative tumour volume described in section

2.3. It is readily apparent that the dose calculated with Monte Carlo is much lower than

the dose calculated within iPlan. For the plans created in iPlan, the V100 values are

95%, 100% and 100% for the 4, 7.5 and the 10 mm cones respectively. In comparison,

the V100 values calculated in MMCTP for the plans calculated with Monte Carlo were

88.8%, 84% and 56% for the 4, 7.5 and 10 mm cones respectively. The largest

difference for the V100 occurred for the plans with the 10 mm diameter cone. This

was larger, in comparison to the 4 and 7.5 mm cones, because the representative target

volume delineated for the 10 mm cone covers a larger area of low density tissue than

the volume delineated for the 4 and 7.5 mm cones. This DVH comparison further

confirms the results obtained for the phantom study in section 3.1 and highlights the



dosimetric limitations of the Clarkson pencil beam used for cones through low density

media.

Figure 7-8 Dose Volume Histogram for the dose distribution calculated in an

anthropomorphic phantom using iPlan and Monte Carlo for a 7.5 mm and 10 mm cone.

7.4 DISCUSSION

In regions where low density media exist, some treatment planning systems over

predict the dose due to the fact that they only account for increased transmission of the

x-rays through the lower density medium18. However, when field size is reduced down

to 1 × 1 cm2 or less, the issue becomes more complicated. When the field size is

reduced, the mean range of secondary electrons in the medium can become larger than

the radius of the field size18, 26, 27. In this particular scenario, the secondary electrons

deposit their energy outside the field and thus contribute less to the dose deposited

within the field28. Therefore, a dual effect takes place as the heterogeneities are

encountered in small x-ray fields. As the density of the medium is increased, the

secondary electron range may be reduced thus contributing to a slight increase in dose

deposited in the medium19. For the low density medium, the radiation dose was

reduced as the field size was reduced down to 4 mm diameter.


radiosurgery 129

Several studies have investigated the accuracy of the pencil beam and

convolution type algorithms for small x-ray fields traversing through low density

tissues. In one study, Jones et al20 study involved comparing inhomogeneity correction

algorithms with Monte Carlo simulations for photon field sizes down to 5 mm

diameter. The algorithms studied were Batho, equivalent path length and convolution

super-position and the benchmark values were obtained from BEAMnrc/DOSXYZnrc

simulations. They found that for the Monte Carlo simulations and the convolution

superposition algorithm, the dose calculated inside the low density inhomogeneity

decreased as the field size was reduced to 5 mm2. In similar work, Carrasco et el18

tested correction based and convolution photon dose calculation algorithms for

radiotherapy for fields down to 1 × 1 cm2 with various x-ray beam energies. They

found that the only algorithm that correctly predicted the penumbra broadening effect

in low density media was the collapsed cone convolution-superposition algorithm. In

water equivalent media, all of the algorithms correctly predicted the dose to within

2%. Our findings are in agreement with both of these studies where the correction

based algorithm in iPlan deviates from the Monte Carlo benchmarked calculations in

low density media.

In a later study, Stathakis et al19 examined the accuracy of Acuros XB and AAA

dose algorithms in Eclipse (Varian Medical Systems, Palo Alto, USA) for small fields

down to 1 × 1 cm2. They considered the calculations performed with

BEAMnrc/DOSXYZnrc as the benchmark in their study. They found that the dose

predicted in low density media, lung equivalent in their case, decreased as the field

size was reduced from 5 × 5 cm2 down to 1 × 1 cm2. Our results are in agreement with

theirs in the sense that our results also show a reduction in predicted dose in low

density media as the field size was decreased down to 4 mm in diameter. Stathakis et

al19 also showed that in high density media, bone equivalent in their case, the dose

predicted increased slightly as the field size was reduced from 5 × 5 cm2 down to 1 ×

1 cm2. This finding was also in agreement with our results where the dose to high

density media increased as the field size was reduced to 4 mm diameter.

If a tumour is located close to or in a region of low/high density media such as

air cavities or bone careful consideration needs to be applied to beam configurations.

Based on the results of this work, beam selection in terms of position and angle much

be optimised in the case of treating very small cancers in the head region. While it is



acknowledged that most of the beam may not pass through low density tissue, some of

the beam may do so. In those cases, the 2D and 3D dose data needs to be carefully

examined with possible large uncertainties for those beams passing through the low

density tissue. In addition, this work has also been the first to perform DVH

calculations for very small tumour volumes with field sizes down to 4 mm diameter.

As shown in the results, there can be large differences in the DVH doses.

Pencil beam algorithm such as the one found in the Brainlab treatment planning

system is well known to be limited when calculating dose in inhomogeneous medium.

There are superior algorithms available in the Brainlab treatment planning system such

as Monte Carlo. However, the only module which calculates dose using the circular

cones is the pencil beam algorithm. It would be beneficial to include Monte Carlo

option for circular cones as well.

Finally, we have demonstrated the feasibility of using a full Monte Carlo dose

calculation for SRS treatments even within patient geometry that contains significant

tissue inhomogeneities. This Monte Carlo system can be used to verity the doses for

cone based treatments and may be used to guide clinical decisions in terms of beam

placement so as to minimise dose inaccuracies due to the limitation of the treatment

planning system algorithm in those specific cases.

7.5 CONCLUSION

This work showed that the iPlan Clarkson algorithm normally used for planning

stereotactic radiosurgery of brain tumours can result in significant difference in

predicted dose when compared with Monte Carlos based simulations for the Brainlab

circular cones of 4, 7.5 and 10 mm in diameter. Extreme care must be used when

placing cone delivered beam arrangements for treatment planning in regions where

there is a large variation in density around the target volume.


Computational resources and services used in this work were provided by the

High Performance Computing and Research Support Unit, Queensland University of

Technology (QUT), Brisbane, Australia.


radiosurgery 131

We also acknowledge the support provided by Dr Andrew Alexander, McGill

University on adapting the MMCTP platform for simulating the Brainlab circular

cones with BEAMnrc/DOSXYZnrc.

We thank Mr Michael O’Connor and Mr Nonga Fangupo, both of Chris O’Brien

Lifehouse, for their assistance in helping to develop the treatment plans in the head

phantom.

7.7 COMPLIANCE WITH ETHICAL STANDARDS

We hereby declare that there is no conflict of interest by any of the authors. This

article does not contain any studies with human participants or animals by any of the

authors.


The authors declare no conflict of interest

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Chapter 8: Discussions and Conclusions 135

Chapter 8: Discussions and Conclusions

8.1 DISCUSSION ON NEW COMMERCIAL DETECTOR – PTW 60019

MICRODIAMOND

The work presented in Chapter 2 was published in 2014 and it was one of the

first studies published1 which evaluated the performance of the then newly released

commercial synthetic diamond detector being the PTW 60019 microDiamond from

PTW (Freiburg, Germany). Following this publication in Medical Physics, there have

been many other publications which have added weight in favour of using the

microDiamond detector for small field measurements2-10 as well as some controversy

to the Monte Carlo simulation of this detectors11, 12.

Andreo et al3 performed Monte Carlo calculations for the PTW 60019

microDiamond to obtain detector-specific output correction factors using geometry

description obtained from the manufacturer blueprint technical drawings. This group

used two Monte Carlo codes for this purpose, one being PENELOPE/penEasy and the

other EGSnrc/egs_chamber. The agreement in calculations using the separate codes

with the same blueprint geometry was within 0.3 % of each other which confirmed the

accuracy in their Monte Carlo simulations. However, the Monte Carlo calculated

output correction factor for the smallest field size was about 3% in disagreement with

the published value in the new TRS-483 Code of Practice, well outside the 90%

confidence limits shown in the TRS-483 Code of Practice. This difference created

some controversy because Andreo et al used the manufacturer blueprint drawings for

their Monte Carlo simulations. How could this be possible? This group then produced

x-ray images of the sensitive volume which after evaluation showed that the effective

area could be potentially different to the blueprint dimensions. After adjustment of the

effective volume in their Monte Carlo calculations to match the dimensions shown in

the x-ray images, they showed that the new Monte Carlo calculated output factor

correction changed by about 3.5% and it now agreed with the published value in the

new TRS-483 Code of Practice. This issue created controversy for the fact that what

the manufacturer provided as blueprint produced unexpected results. In the conclusion

of their work, this group then warned caution to physicists about using blueprint

drawings from the manufacturer3.

136 Chapter 8: Discussions and Conclusions

In response to this paper, Marinelli et al4 published a paper on the reproducibility

of the fabrication process for the PTW 60019 microDiamond detector. This group

estimated the average active surface area by recording 2D maps under scanned soft x-

ray microbeam irradiation. They also used capacitance measurements and alpha

particle detection experiments. Their results found that the average active surface area

diameter indeed agreed with that provided by the manufacturer. They found that no

contributions were observed from the housing or encapsulation materials of the device.

This group advised that there should be further investigation on the role of volume

average effects and other perturbation effects separately in order to clarify the

discrepancy.

After Marinelli published their article, there was a reply from Andreo et al12

which, in short, highlighted the fact that the microDiamond provided in the

manufacturer’s blueprint was not the real microDiamond when inspected by x-rays.

And to obtain good agreement between Monte Carlo simulation and measurement

required the simulation of an “effective microDiamond” as opposed to the real device.

This healthy debate provided a boost to the attention given to the new microDiamond

detector.

Further work was published by a number of groups including Larraga et al2,

Mancosu et al13, Masi et al7 and De Coste et al8. All these authors published results

which supported the use of the PTW 60019 microDiamond detector for use in

measurements of small x-ray fields down to 5 × 5 cm2. Thus, confirming the finding

in this thesis on the suitability of this detector.

In this thesis, as shown in Chapter 2, relative dosimetry measurements with the

microDiamond were performed for small x-ray fields down to 4 mm diameter. This

dosimetry data was compared with both BEAMnrc Monte Carlo calculations as well

as dosimetry data measured using diodes detectors that were considered acceptable at

that time. It was shown that the microDiamond detector is a suitable candidate for

small field dosimetry for the Brainlab circular cones down to 4 mm diameter. It was

found that the percentage depth doses measured with the microDiamond detector for

the 4, 7.5, 10 and 30 mm agreed to within 1.5% between two other commercial diode

detectors namely the IBA SFD stereotactic diode and the PTW 60012 Diode E

detector. For measurements of cross profiles and penumbra in water at depth, it was


found that sharper penumbra results were obtained, for all the cone diameters, when

the microDiamond was oriented in a perpendicular direction to the x-ray beam.

The maximum correction factor, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , derived for this detector under these

conditions for the 4, 7.5, 10, 20 and 30 mm cone was 1.027. The correction factors

provided in the new CoP TRS-48310 are values for different geometric set up where

the detector is at 10 cm depth thus further away from the source and where the volume

averaging effect is reduced. Thus, the results in this paper should be used only for this

geometry, as required by the treatment planning system from Brainlab as it is not

considered in the Code of Practice.

The suitability of the PTW microDiamond detector for measurements in small

field dosimetry was further confirmed by work subsequent to the publication of the

new CoP TRS-483. For example, it has been evaluated as a reference detector in Italy

for small field dosimetry audits within their primary standards laboratory8. They

concluded that, in principle, accurate reference dosimetry is feasible by using the

microDiamond dosimeter for field sizes down to 5 mm. In addition, in 2017, the

Australian Radiation Protection and Nuclear Safety Agency (ARPANSA), evaluated

a number of detectors for measurement of a 5 mm diameter cone14 including the PTW

60019 microDiamond detector. Subsequently, the Australian Clinical Dosimetry

Service (ACDS) run by ARPANSA, uses the PTW 60019 microDiamond to perform

small field output factors measurements for Level 1b Dosimetry Audits for SBRT and

SRS which involve very small fields.

8.2 DISCUSSION ON MONTE CARLO MODELLING OF GAFCHROMIC

EBT3 FILM

In Chapter 3, output correction factors, 𝑘𝑄𝑐𝑙𝑖𝑛,𝑄𝑚𝑠𝑟

𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were calculated for

Gafchromic EBT3 film using Monte Carlo simulations. These factors were determined

for a Novalis Trilogy linear accelerator equipped with Brainlab circular cones with

diameters of 4.0 to 30.0 mm. The Monte Carlo code used to simulate the x-ray beam

using the Brainlab code was BEAMnrc. And the Monte Carlo code to simulate

Gafchromic EBT3 film was DOSXYZnrc.

The results showed that for circular cones from 5 to 30 mm in diameter at an

SSD of 100 cm and 10.0 cm depth confirmed that the output correction factors,



𝑓𝑐𝑙𝑖𝑛,𝑓𝑚𝑠𝑟 , were equal or less than 0.4%. It was shown that for fields with a diameter

less than 5 mm the output correction factor was equal or less than 1%. This study

further supports the assumption that radiochromic film is a correction less dosimeter

and that if used appropriately it can be used as a reference detector for small field

dosimetry when these output correction factors are taken into account.

8.3 DISCUSSION ON EXTRAPOLATION TECHNIQUE FOR

GAFCHROMIC EBT3 FILM

An experimental extrapolation technique using Gafchromic EBT3 film for

measurement of relative output factors15 was presented in Chapter 4. The relative

output factor for Brainlab circular cones of 4 to 30 mm2 diameter were determined by

reducing the circular region of interest (ROI) and extrapolating to zero area to remove

effects of volume averaging. The results found that a high scanning resolution of 1200

dpi was required in order to obtain enough data points and to minimise noise in the

signal. It was shown that the extrapolated relative output factor for the 4 mm diameter

cone was 0.651 but when the diameter of analysis area was varied from 0.5 to 1.0 mm

diameter the corresponding relative output factors were 0.639 and 0.633 which

corresponded to a relative output factor change of 1.8% and 2.8%. It was concluded

that for very small fields such as a 4 mm diameter cone a measurable difference can

be seen in the relative output factor based on the ROI and the size of the area of

analysis. This paper concluded that it was recommended to scan the Gafchromic EBT3

film at a resolution of 1200 dpi for circular cones with sizes of less than 7.5 mm in

diameter to utilise the extrapolation technique to obtain the relative output factor. This

technique is purely experimental and very useful for departments where Monte Carlo

simulation facilities are not available.

8.4 DISCUSSION ON EXTRAPOLATION TECHNIQUE FOR OSLD

DETECTORS

In Chapter 5, a novel extrapolation method using OSLDs was presented. This

method involved varying the size of optically stimulated luminescent dosimeters

(OSLDs) to determine relative output factors for field diameters ranging from 4 to 30

mm as defined by the Brainlab SRS cones. Water droplets were used to remove air

gaps located around the OSLD detectors thus minimising density variation effects. The

results for the 4 mm diameter cone showed that the relative output factor with this


technique was 0.660 compared to 0.661 ± 0.01 and 0.651 ± 0.018 for the PTW 600019

microDiamond detector and Gafchromic EBT3 film respectively. It was concluded

that OSLDs can be capable of producing results of similar accuracy to the

microDiamond detector and to Gafchromic EBT4 film.

8.5 DISCUSSION ON SKIN/SURFACE DOSE FOR BRAINLAB

CIRCULAR CONES

The work presented in Chapter 6 involved a comparison of skin dose calculated

through Monte Carlo methods and through measurements with Gafchromic EBT3

film16. Skin dose is a potential clinical concern in SRS treatments, due to the high

single-fraction doses delivered, and there is a growing interest in the incidence of skin

toxicity associated with stereotactic body radiotherapy (SBRT or SABR)17. Prior to

this work there was limited data in terms of the skin dose for the very small field sizes

used in these treatments. Relative skin doses were determined for Brainlab circular

cones ranging in sizes of 4 to 30 mm diameter. Monte Carlo calculations were

performed using the BEAMnrc code with a model of the Novalis Trilogy linear

accelerator and the Brainlab circular collimators. Surface doses were calculated at the

ICRP skin dose depth of 70 m using the 6 MV SRS x-ray beam. These calculate doses

were in good agreement to the measured data with an agreement of better than 2%.

This work indicates that Gafchromic EBT3 film can be used for accurate surface dose

measurements in clinical situations where Monte Carlo simulations might not

available for stereotactic radiosurgery beams.

8.6 DISCUSSION ON INHOMOGENENITY CORRECTION ON SMALL

FIELDS PRODUCED BY BRAINLAB CIRCULAR CONES AND

IMPLEMENTATION OF MMCTP FOR MONTE CARLO BASED

INDEPENDENT CHECKS

The remaining chapter in this thesis has applied Monte Carlo techniques to verify

the dose to tissue inhomogeneities for very small fields. The first part of the study

involved simple phantom geometries with inhomogeneities added on. This work

proved that the simple dose calculation algorithm used for circular cones in the

Brainlab system may provide large deviations in tissue inhomogeneneities and as such

care must be taken. This has clinical application when routine treatment planning

system calculations that make use of circular cones pass through significant tissue

inhomogeneity. The results can provide data for the physicist and may guide


adjustments to the planning process when there are inhomogeneities present. This

chapter also shows how the McGill package, MMCTP, can be used to perform Monte

Carlo based independent checks for iPlan treatment plans for patient treatments using

circular cones. This is useful as there are no commercial packages that can convert

DICOM files from the planning systems to EGSnrc files for Monte Carlo calculations.

Monte Carlo based system are considered gold standard especially for very small

fields. This paper showed that Monte Carlo based calculations can provide improved

penumbra results which can improve the Dose Gradient analysis of treatment plans.

8.7 CLINICAL IMPLICATIONS OF THE WORK

The work in this thesis can be combined to obtain improvement in small field

dosimetry in a clinical radiotherapy department where stereotactic treatments take

place. For example, the output correction factors presented in Chapter 2 for the

microDiamond detector for measuring output factors can be implemented when

commissioning a system with Brainlab circular cones. The data measured for the

Treatment Planning System can performed using this detector and at present the user

can even implement the IAEA TRS-483 Code of Practice for this purpose. In this

thesis, the Brainlab planning system was used. However, this methodology can be

applied to any of the other commercial treatment planning systems available in the

market such as CyberKnife and GammaKnife systems.

The techniques of chapter 3, 4 and 5 can be combined to obtain improved

accuracy for very small fields in the range of 5 to 10 mm in diameter. The improvement

in the use of EBT3 film for performing measurements for point dose measurements as

well as 2 D dose maps is of paramount importance in the clinic. The methodology

presented in these chapters is an improvement over current clinical techniques.

The work presented in Chapter 6 can be used to test a treatment planning system

for the assessment of skin dosimetry. Treatment planning systems do not predict the

skin dose very accurately. This issues is very well known with planning systems. The

results presented in this chapter can be used to assess the accuracy when

commissioning a system such as the Brainlab system equipped with circular cones.

The work in Chapter 7 is an example of how everything presented in this thesis

can be combined together. If the measurement techniques shown in chapters 2 to 6 can

be applied when taking measurements for data for the planning system then there will


be an improvement in the accuracy of the whole process. The same techniques can

then be applied when verification of the planning system takes place during

commissioning and for subsequent Quality Assurance of patient-specific plans.

8.8 CONCLUSIONS

Accurate dosimetry is critical for stereotactic radiosurgery where very small

fields are used to treat brain tumours. Field sizes are usually less than 15 mm in

diameter and this places these fields in the region of electronic disequilibrium. A wide

range of knowledge and skills is required to encompass the great challenge in this area.

The aim of the work presented in this thesis was to contribute to key aspects in this

area and provided guidance to solve some of the challenges. In conclusion, the work

in this thesis has proved that

• the new PTW 60019 microDiamond detector is suitable for taking

measurements in small x-ray fields below 30 mm diameter down to 4

mm diameter.

• Monte Carlo modelling of Gafchromic EBT3 film has indeed proved that

EBT3 is water equivalence for measurements in small fields down to 4

mm in diameter as has been assumed in the literature.

• The implementation of a novel extrapolation technique using

Gafchromic EBT3 film to measure the relative output factor for the 4 mm

diameter Brainlab cone is valid and can be implemented in the clinic.

• It is possible to re-design and use existing commercial OSLDs for

obtaining the relative output factor for a 4 mm diameter Brainlab cone

using the extrapolation technique.

• The surface or skin dose can be obtained using Monte Carlo methods for

the Brainlab circular cones and can be used to commission new treatment

planning system for stereotactic radiosurgery and

• Monte Carlo methods can be used to verify the dose distribution

produced by the current dose calculation algorithm in the Brainlab

system used for circular cones.


8.9 FUTURE WORK IN SMALL FIELD DOSIMETRY FOR

RADIOSURGERY

There are still many challenges in small x-ray field dosimetry that require further

investigation by the medical physics community worldwide. There will be new

radiation dosimeters developed and released for use in small fields. Their use in small

x-ray field measurements will require additional investigation into whether they may

or may not require detector correction factors for use with the IAEA TRS-483 Code

of Practice. The work presented in this thesis, provides a framework by which these

correction factors can be readily determined by using Gafchromic EBT3 film and

Monte Carlo calculations.

Further work is also required in Monte Carlo models for very small fields. The

IAEA TRS-483 Code of Practice highlighted the fact that the spectrum produced by

Monte Carlo models might not be the same and it would be beneficial to unify all

Monte Carlo based models so that universal unification be obtained especially when

lateral charged particle equilibrium is lost. In fact, one good suggestion would be to

concentrate effort in modelling fields below 15 mm in width. Currently most Monte

Carlo models are optimised for dosimetry for reference fields which are usually greater

than 6 cm in diameter.

The formation of a new Task Group, similar to Task Group 10518, to unify Monte

Carlo models for different technologies available (GammaKnife, CyberKnife,

Tomotherapy, TrueBeam and Elekta linacs) with emphasis on fields smaller than 15

mm would be advisable. The work of this task group would be similar to other tasks

groups from the past such as IPEM8119 and many others over the years. There are

currently a few Monte Carlo codes available as freeware such as EGSnrc, Penelope,

Fluka and GEANT4. If it would be possible to set benchmarks so that all codes can

meet a given criteria for small field dosimetry perhaps using the 15 mm diameter field

as the benchmark, then this could lead to a more consistent set of results.

More work can also be done improving the accuracy of the dosimetric

performance of Gafchromic EBT3 film. At the moment Gafchromic EBT3 film can

produce considerable noise due to its granular structure, however, if this could be

improved so that a more consistent signal could be obtained then this would also help

to facilitate more consistent results from one film batch to another.


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advances in very small x ray field dosimetry for … · mm circular fields for which limited...

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