Hyperthermia Mediated Drug Delivery using Thermosensitive ...The clinical efficacy of chemotherapy...

Hyperthermia Mediated Drug Delivery using Thermosensitive Liposomes and

MRI-Controlled Focused Ultrasound

by

Robert Michael Staruch

A thesis submitted in conformity with the requirements for the degree of Doctor of Philosophy

Department of Medical Biophysics University of Toronto

© Copyright by Robert Michael Staruch 2013

iii

Abstract

Hyperthermia Mediated Drug Delivery using Thermosensitive Liposomes and MRI-Controlled Focused Ultrasound

Robert Michael Staruch

Doctor of Philosophy

Department of Medical Biophysics

University of Toronto

2013

The clinical efficacy of chemotherapy in solid tumours is limited by systemic toxicity and

the inability to deliver a cytotoxic concentration of anticancer drugs to all tumour cells.

Temperature sensitive drug carriers provide a mechanism for triggering the rapid

release of chemotherapeutic agents in a targeted region. Thermally mediated drug release

also leverages the ability of hyperthermia to increase tumour blood flow, vessel

permeability, and drug cytotoxicity. Drug release from thermosensitive liposome drug

carriers in the tumour vasculature serves as a continuous intravascular infusion of free

drug originating at the tumour site. However, localized drug release requires precise

heating to improve drug delivery and efficacy in tumours while minimizing drug

exposure in normal tissue.

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Focused ultrasound can noninvasively heat millimeter-sized regions deep within

the body, and can be combined with MR thermometry for precise temperature control.

This thesis describes the development of strategies to achieve localized hyperthermia

using MRI-controlled focused ultrasound, for the purpose of image-guided heat-triggered

drug release from thermosensitive drug carriers.

First, a preclinical MRI-controlled focused ultrasound system was developed as a

platform for studies of controlled hyperthermia and drug delivery in rabbits. The

feasibility of using ultrasound hyperthermia to achieve localized doxorubicin release

from thermosensitive liposomes was demonstrated in normal rabbit muscle. Second,

strategies were described for using MR thermometry to control ultrasound heating at a

muscle-bone interface based on MR temperature measurements in adjacent soft tissue,

demonstrating localized drug delivery in adjacent muscle and bone marrow. Third,

fluorescence microscopy was employed to demonstrate that increased overall drug

accumulation in rabbit VX2 tumours corresponds to high levels of bioavailable drug

reaching their active site in the nuclei of tumour cells.

The results of this thesis demonstrate that image-guided drug delivery using

thermosensitive liposomes and MRI-controlled focused ultrasound hyperthermia can be

used to noninvasively achieve precisely localized drug deposition in soft tissue, at bone

interfaces, and in solid tumours. Clinical application of this work could provide a

noninvasive means of enhancing chemotherapy in a variety of solid tumours.

v

Acknowledgements

I owe a great deal of thanks to my supervisor, Dr. Rajiv Chopra, for his mentorship,

encouragement, and support. His energy and enthusiasm inspired me to keep moving forward,

and I thank him for letting me take the time to develop a broad range of experience. I'm fortunate

to have witnessed the early stages of his successful academic career: the creation of an academic

research lab, two companies, and two clinical trials. Thank you for keeping me in the loop on

these developments, I hope I can apply some of those lessons going forward. Thanks also for

letting me take part in the prostate human study; seeing the potential of this technology to

improve people's lives makes all of the work worthwhile. Equal thanks go to Dr. Kullervo

Hynynen, for sharing his incomparable experience, for continuing to support my expensive

experiments, and for teaching me how to focus on the path to results. I'm also extremely grateful

to Dr. Ian Tannock, for his patience, for his advice (scientific, medical, and grammatical), and

for allowing me to learn from his staff how to use their microscopy techniques.

The experiments that make up my thesis would not have been possible without many

skilled and dedicated people working behind the scenes. I'd especially like to thank Alex Garces

and Shawna Rideout, who’ve put up with years of nagging requests and ridiculous hours, where

"Oh yeah, we'll be done by nine" really meant recovering rabbits until 3 a.m. I'm greatly indebted

to the skill and creativity of the machinists in our group: Anthony Chau, the mastermind of the

focused ultrasound system used in my experiments, and Moe Kazem, who's always made time to

design perfect parts for my setup. I'd also like to thank Milan Ganguly, Miria Bartolini, and

Jasdeep Saggar for helping this hopeless engineer learn the black art of fluorescence microscopy.

These studies could not have been undertaken without funding from the Terry Fox Foundation,

the Ontario Institute for Cancer Research, and NSERC Canada Graduate Scholarships. I’m also

indebted to Celsion for generously providing me with thermosensitive liposomal doxorubicin.

I wouldn't have been able to perform or even understand these experiments were it not for

the mentorship of many intelligent, generous people at Sunnybrook. Dave Goertz, Sam Pichardo,

Kee Tang, Kevan Anderson, Junho Song, Sam Gunaseelan, Ben Lucht, Meaghan O'Reilly,

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Aaron Boyes, and Charles Mougenot, thank you for giving me so much of your time to answer

my naive, often repetitive questions.

It's been a privilege to work with several generations of staff and students in the Focused

Ultrasound Lab. My heartfelt thanks go to those who shared with me the special bond of being

rowmates: Ian Pang, Alison "Dr. Alleycat" Burgess, and Mathew Carias (Vanier Scholar). Arvin

Arani, it's been great to have you as a teammate on the original Team Chopra. It would be remiss

if I didn’t thank Ali Yousefi, Alex Klotz, Dan Pajek, Ryan Jones, Nick Ellens, and Leila Shaffaf

for always being in the mood for a coffee, and of course, Cameron Wright, for literally saving

me from drowning. To everyone in C7, our chats at the front of the room were a highlight of

day-to-day life in the lab. I wish you all continued success.

I'd also like to thank a few jarns for their friendship over the years: my roommates at

Davisville, Rahul Sarkar and Brandon Helfield, and the original sickest guy ever, Dr. Mark

Chiew. Thank you for your support in tough times, and for your patience when I was distracted

by work or other pursuits. It has been an honour serving with you on the bridge.

I wouldn't be here without my family, and I am sincerely thankful to them. My sister

Andrea, for her support and friendship, and my parents, for all the things they gave up in order to

put us and our education first. Mom and Dad, thank you for getting me through my experience

with cancer. With your help, it became a turning point in my life, and led me into what has

become my life's work.

Finally, Michelle, who throughout the course of my PhD has suffered more than anyone.

Thank you for tolerating entire weekends at the library, for driving me home from the lab in the

middle of the night, and for supporting me when all hope was lost. I will always be grateful for

the sacrifices you've made to help me reach my goals, and I look forward to this next stage of life

together.

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Table of Contents

Chapter 1 Introduction and Background ......................................................................................... 1

1.1 Barriers to Drug Delivery in Solid Tumours .................................................................... 1

1.1.1 Cancer ....................................................................................................................... 1

1.1.2 Cancer Chemotherapy ............................................................................................... 2

1.1.3 Tumour microenvironment ....................................................................................... 4

1.1.4 Doxorubicin .............................................................................................................. 7

1.2 Hyperthermia .................................................................................................................. 13

1.2.1 Cytotoxic effects ..................................................................................................... 13

1.2.2 Chemosensitization ................................................................................................. 14

1.2.3 Blood flow .............................................................................................................. 15

1.2.4 Vessel Permeability ................................................................................................ 17

1.3 Hyperthermia mediated drug delivery ............................................................................ 17

1.3.1 Liposomes ............................................................................................................... 17

1.3.2 Thermosensitive liposomes ..................................................................................... 20

1.3.3 Clinical Experience ................................................................................................. 23

1.4 MRI-Controlled Focused Ultrasound Hyperthermia ...................................................... 25

1.4.1 Clinical Hyperthermia and Thermal Therapy Techniques ...................................... 25

1.4.2 Focused Ultrasound ................................................................................................ 27

1.4.3 MR Image Guidance and Thermometry ................................................................. 30

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1.4.3.1 Magnetic Resonance Imaging ......................................................................... 30

1.4.3.3 MRI Thermometry ........................................................................................... 32

1.4.5 MRI-Controlled Focused Ultrasound ..................................................................... 35

1.4.5.1 Motivation for feedback control ...................................................................... 35

1.4.5.2 Scanned focused ultrasound hyperthermia ...................................................... 37

1.4.5.3 Proportional-integral-derivative control of focused ultrasound thermal therapy

......................................................................................................................... 39

1.5 Thermally mediated drug delivery using MRI-controlled focused ultrasound .............. 41

1.6 Specific Aims ................................................................................................................. 43

Chapter 2 Localized drug release using MRI-controlled focused ultrasound hyperthermia ........ 47

2.1 Introduction .................................................................................................................... 47

2.2 Materials and Methods ................................................................................................... 49

2.2.1 MRI-controlled focused ultrasound system ............................................................ 49

2.2.2 Animals ................................................................................................................... 50

2.2.3 MR imaging ............................................................................................................ 53

2.2.4 Closed-loop feedback control ................................................................................. 54

2.2.5 Numerical Simulations ............................................................................................ 56

2.2.6 Sonication, drug administration and tissue harvesting ........................................... 56

2.2.7 Analysis of drug concentration and release ............................................................ 57

2.3 Results ............................................................................................................................ 58

2.3.1 Simulations and in vitro temperature control .......................................................... 58

ix

2.3.2 MRI-guided focused ultrasound hyperthermia in vivo ........................................... 59

2.3.3 Hyperthermia mediated drug delivery .................................................................... 65

2.4 Discussion ...................................................................................................................... 67

2.4.1 MRI-controlled focused ultrasound hyperthermia .................................................. 67

2.4.2 Sources of error in MRI thermometry .................................................................... 68

2.4.3 Hyperthermia mediated drug delivery .................................................................... 68

2.4.4 Limitations and sources of variability in drug delivery .......................................... 70

2.5 Conclusions .................................................................................................................... 71

Chapter 3 MRI-controlled focused ultrasound hyperthermia for targeted drug delivery in bone: in

vivo results .................................................................................................................................... 73

3.1 Introduction .................................................................................................................... 73


3.2.1 Animals ................................................................................................................... 74

3.2.2 Focused ultrasound system ..................................................................................... 75


3.2.4 Simulations of temperature elevations in bone ....................................................... 79

3.2.5 Drug administration and analysis of drug concentrations in tissue ........................ 79

3.2.6 Statistical Analysis .................................................................................................. 81

3.3 Results ............................................................................................................................ 81


3.3.2 Numerical simulations of temperature elevations in bone during MRI-controlled

hyperthermia ......................................................................................................................... 82

x

3.3.3 Doxorubicin concentrations in bone marrow and muscle ....................................... 87

3.4 Discussion ...................................................................................................................... 90

3.4.1 Hyperthermia and drug delivery in bone ................................................................ 90

3.4.2 Practical Applications ............................................................................................. 91

Chapter 4 Enhanced drug delivery in rabbit VX2 tumours using thermosensitive liposomes and

MRI-controlled focused ultrasound hyperthermia ........................................................................ 93

4.1 Introduction .................................................................................................................... 93


4.2.1 Animals and VX2 tumours ..................................................................................... 94


4.2.3 Drug concentration in unheated tissue and VX2 tumours ...................................... 96

4.2.4 Drug distribution in the tumour microenvironment ................................................ 97

4.3 Results ............................................................................................................................ 98

4.3.1 Animals and VX2 tumours ..................................................................................... 98

4.3.2 MRI-controlled focused ultrasound hyperthermia in VX2 tumours ..................... 100

4.3.3 Drug deposition in VX2 tumours .......................................................................... 100

4.3.4 Increased delivery of bioavailable drug in tumour cells ....................................... 104

4.4 Discussion .................................................................................................................... 108

4.5 Conclusion .................................................................................................................... 112

Chapter 5 Conclusions and Future Work .................................................................................... 113

5.1 Summary of findings .................................................................................................... 113

5.2 Limitations ................................................................................................................... 116

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5.2.1 Quantification of doxorubicin fluorescence .......................................................... 116

5.2.1.1 Doxorubicin fluorometry in homogenized tissue samples ............................ 117

5.2.1.2 Fluorescence microscopy in frozen tumour sections ..................................... 119

5.3 Further studies: The effect of localized drug delivery on antitumour efficacy ............ 120

5.3.1 Study motivation and design ................................................................................. 120

5.3.2 Doxorubicin antitumour efficacy and toxicity ...................................................... 121

5.3.3 Using a clinical HIFU system for MRI-controlled hyperthermia ......................... 123

5.4.4 Early results .......................................................................................................... 125

5.4 Future directions and clinical applications ................................................................... 127

5.4.1 Thermally-mediated drug delivery using other therapeutic agents and delivery

vehicles ............................................................................................................................... 127

5.4.2 Imageable nanoparticle drug carriers .................................................................... 129

5.4.3 Clinical applications for thermally mediated drug delivery ................................. 130

References ................................................................................................................................... 133

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List of Tables

Table 2.1: Summary of in vivo MRI-controlled FUS experiments for hyperthermia-mediated

drug delivery. ................................................................................................................................ 63

Table 3.1: MR imaging parameters. ......................................................................................... 77

Table 3.2: Acoustic and thermal properties for numerical simulations of scanned focused

ultrasound heating in bone. ........................................................................................................... 80

Table 3.3: Summary of in vivo MRI-controlled focused ultrasound hyperthermia experiments

with the focus set at two different offsets from a muscle-bone interface in rabbit thigh. ............ 84

Table 3.4: Numerical simulations of MRI-controlled focused ultrasound hyperthermia in

bone, with the focus set at two different offsets from the muscle-bone interface. ....................... 87

Table 3.5: Doxorubicin concentrations measured by fluorescence intensity in tissue samples

harvested from heated and unheated regions of rabbit thigh muscle and bone marrow. .............. 89

Table 4.1: Temperatures measured noninvasively with MR thermometry during MRI-

controlled focused ultrasound hyperthermia in rabbit VX2 tumours. ........................................ 101

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List of Figures

Figure 1.1: Barriers to drug delivery in the solid tumour microenvironment. ............................ 4

Figure 1.2: Fluorescence spectrum of doxorubicin. ................................................................ 8

Figure 1.3: Cellular uptake and cytotoxic effects of doxorubicin in human lung cancer cells

[51]. ................................................................................................................................. 10

Figure 1.4: Doxorubicin distribution in tumour and normal tissue in MDA-MB-231 tumour

bearing mice at 10 minutes after injection. ................................................................................... 11

Figure 1.5: Left: Microregional distribution of doxorubicin fluorescence in a cryosection of a

locally advanced breast cancer biopsy at 24 h after i.v. injection of doxorubicin. ....................... 12

Figure 1.6: Effect of hyperthermia on blood flow in normal granulating tissue (solid) and VX2

tumours (dashed) in rabbit ear window preparation. .................................................................... 16

Figure 1.7: Structure and temperature-sensitivity of liposomal drug carriers. ...................... 22

Figure 1.8: Localized intravascular drug release from thermosensitive liposomes. ................. 24

Figure 1.9: Strategies for using thermal conduction to heat large tissue volumes using focused

ultrasound. ................................................................................................................................. 28

Figure 1.10: Temperature measurements made with MRI using the proton resonance

frequency shift phase-difference technique. ................................................................................. 34

Figure 1.11: Localized drug delivery using thermosensitive liposomes and MRI-controlled

focused ultrasound. ....................................................................................................................... 42

Figure 2.1: Experimental setup for in vivo MRI-controlled focused ultrasound hyperthermia. 51

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Figure 2.2: Simulations (A,C,E) and in vitro (B,D,F) testing of proportional-integral MRI

temperature control for mechanically-scanned focused ultrasound hyperthermia. ...................... 60

Figure 2.3: Evolution of temperature distribution during MRI-controlled focused ultrasound

heating of a 10 mm diameter region. ............................................................................................ 61

Figure 2.4: Example of temporal and spatial temperature control over a 10 mm diameter

circular area using MRI-controlled focused ultrasound hyperthermia. ........................................ 62

Figure 2.5: Tissue damage observed using T2-weighted and contrast-enhanced T1-weighted

imaging following thermal coagulation in rabbit thigh. ............................................................... 64

Figure 2.6: Doxorubicin concentrations measured by fluorescence intensity in tissue samples

harvested from heated and unheated regions of thigh muscle in rabbits receiving controlled

hyperthermia (n = 6) or thermal coagulation (n = 4). ................................................................... 66


from rabbit #9, harvested at 0, 5, 10, and 20 mm away from the center of the heated region, as

well as from the unheated contralateral thigh. .............................................................................. 67

Figure 3.1: Experimental setup for MRI-controlled focused ultrasound heating in bone. ........ 76

Figure 3.2: Snapshot of temperature distribution from controlled hyperthermia at muscle-bone

interface in rabbit thigh using MRI-controlled focused ultrasound. ............................................. 82

Figure 3.3: Temporal evolution of controlled hyperthermia at muscle-bone interface in rabbit

thigh using MRI-controlled focused ultrasound, for the same experiment shown in Figure 3.2. 83

Figure 3.4: Radial distribution of thermal dose and temporal mean ± SD steady-state

temperature in A) control plane and B) bone plane, for the same experiment shown in Figures 3.2

and 3.3. ................................................................................................................................. 83

Figure 3.5: Numerical simulations of MRI-controlled focused ultrasound hyperthermia at a

muscle-bone interface. ................................................................................................................. 85

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Figure 3.6: Numerically simulated average steady-state temperature vs. radius for two control

scenarios: control plane set 0 mm from bone interface (empty markers), and control plane offset

10 mm from the bone interface (filled). ........................................................................................ 86


harvested from heated and unheated regions of rabbit thigh muscle and bone marrow. .............. 88

Figure 4.1: MRI-controlled focused ultrasound hyperthermia in rabbit VX2 tumour. ............. 99

Figure 4.2: MRI detection of tissue damage in tumour-bearing rabbit treated with hyperthermia

and thermosensitive liposomal doxorubicin. .............................................................................. 102

Figure 4.3: Plasma and tissue doxorubicin concentrations in tumour-bearing rabbits. ........... 103

Figure 4.4: Microregional distribution of doxorubicin in heated and unheated tumours. ....... 105

Figure 4.5: Heterogeneity of microregional distribution of doxorubicin in regions with varying

tumour vascularity. ..................................................................................................................... 106

Figure 4.6: Accumulation of doxorubicin in the cells of heated and unheated tumours

following administration of thermosensitive liposomal doxorubicin. ........................................ 107

Figure 4.7: Spatial distribution of doxorubicin fluorescence with respect to vessel density and

vessel location in heated and unheated tumours of rabbits administered thermosensitive

liposomal doxorubicin.. .............................................................................................................. 108

Figure 5.1: Sonication and MR thermometry setup using Philips MR-HIFU system. ........ 124

Figure 5.2: Treatment planning and MRI-controlled hyperthermia in a rabbit VX2 tumour

using clinical MR-HIFU system. ................................................................................................ 126

Figure 5.3: Rabbit VX2 tumour response following thermally-mediated drug delivery using

thermosensitive liposomal doxorubicin (TLD) and MRI-controlled focused ultrasound. ......... 126

1

Chapter 1 Introduction and Background

1.1 Barriers to Drug Delivery in Solid Tumours

1.1.1 Cancer

Based on 2011 incidence and mortality rates, 40% of women and 45% of men in Canada will

develop cancer during their lifetimes, and one in four Canadians will die from the disease; solid

tumours represent 85% of all cancers, and 90% of cancer deaths [1].

Cancer is a disease of dysfunctional cell growth that arises from the slow, inefficient

accumulation of genetic mutations that confer survival advantage to increasingly proliferative

abnormal cell types. A tumour is initiated when a single normal cell acquires a mutation that

allows it to outgrow adjacent normal cells. Other mutations and environmental factors increase

the occurrence of genetic alterations in this expanding population of cells, giving rise to

heterogeneous tumour subpopulations. Nearly all of these variants are eliminated because of

metabolic disadvantage or immunologic destruction, but occasionally one has an additional

selective advantage, allowing this mutant to outgrow other tumour cells and become the

precursor of a new predominant subpopulation. Over the course of many years, several

successive waves of proliferation of accidentally advantageous genetic alterations occur in

response to specific selection pressures presented by the tumour’s environment, allowing

tumours to commandeer several of the cell’s essential mechanisms. Along this slow march from

benign mass towards malignancy, tumour cells develop self-sufficiency in growth signals by

activating oncogenes to cause pathological mitosis, acquire insensitivity to anti-growth signals

by inactivating tumour suppressor genes, lose the ability to undergo programmed cell death by

suppressing apoptosis genes and pathways, develop limitless replicative potential by activating

telomerases to retain immortality after many generations, overcome limitations on tumour

2

growth by recruiting new blood vessels through sustained angiogenesis, and eventually gain

mobility and the ability to invade tissue enabling further growth and metastasis to other sites in

the body [2]. The frequency of genetic alterations accelerates as the neoplasm develops; when

faced by new evolutionary challenges such as therapeutic interventions, most cells die, leaving

behind those that by chance have yet another mutation allowing them to escape death and rapidly

repopulate the tumour.

Treatment of many types of cancer focuses on local or regional intervention such as

surgery or radiotherapy to remove the primary tumour mass. Chemotherapy is often administered

after local treatment as an adjuvant therapy, preventing disease recurrence by eliminating cancer

cells at the treatment margins and eradicating micrometastases. It can also be used before local

intervention in a neoadjuvant role, decreasing the size of the primary tumour and the probability

of metastasis. Despite advances in understanding of the mechanisms responsible for

carcinogenesis, surgery and radiotherapy remain the primary therapies for solid tumours, and the

majority of anticancer agents used in the clinic are active against rapidly proliferating cells, with

little specificity for tumours.

1.1.2 Cancer Chemotherapy

Most cytotoxic chemotherapy agents used in the clinic are designed to kill or destroy the

proliferative capacity of rapidly cycling cells, and therefore have limited specificity towards

neoplastic cells. As a result, these anticancer drugs also cause toxicity to rapidly proliferating

normal tissues in the bone marrow and intestinal mucosa. To permit recovery of blood counts

and prevent infection and bleeding, these agents are applied over several treatment cycles, whose

dose and frequency are limited by the toxicity inflicted upon rapidly cycling healthy tissues.

Various classes of chemotherapeutics are defined according to the specific mechanisms by which

they kill cells. Cytotoxic drugs that act by different mechanisms and have different dose-limiting

toxicities can be safely combined in the clinic at or near their maximum tolerated dose to achieve

additive antitumour effect with less than additive normal tissue toxicity.

3

Many human cancers either do not respond to chemotherapy, or acquire resistance during

the course of therapy through a variety of mechanisms occurring at the molecular, cellular, and

microenvironmental level of the tumour. Cellular resistance can arise from decreased drug

uptake through under-expression or functional genetic mutations to cell surface receptors or

transporter proteins responsible for facilitated diffusion and active transport of specific drugs into

cells [3, 4]. Conversely, over-expression of trans-membrane exporter proteins such as multidrug

resistance protein (MRP) and p-glycoprotein (PgP) causes resistance through increased removal

of many natural chemotherapeutics agents from cells [3, 5]. Clinically, efforts to retain

chemotherapeutics within the cell by pharmacologically blocking the action of PgP or MRP have

shown only modest effects in reversing this form of multi-drug resistance [6, 7]. Cells can also

gain resistance through molecular mechanisms, by up-regulating enzymes involved in metabolic

transformation of anticancer drugs, by over-expressing anti-apoptotic signals, or through

mutations to or changes in expression of proteins targeted by cytotoxic agents, thereby

decreasing the cytotoxic activity of drugs despite unaltered intracellular concentration [4].

Recently, increased knowledge in the mechanisms responsible for carcinogenesis have

enabled the intensive development of molecular agents that target the molecular changes and

altered signaling pathways of cancer cells [8-10]. Monoclonal antibodies can be used to bind a

specific circulating ligand, thus interfering with its activity. One such example is bevacizumab,

which binds vascular endothelial growth factor to inhibit angiogenesis [11]. Antibodies can also

target mutated cell surface receptors involved in malignant cellular pathways, such as human

epidermal growth factor receptors (HER) 1 and 2 (e.g. cetuximab and trastuzumab, respectively).

Binding to these receptors inhibits their ability to activate uncontrolled signaling of intracellular

pathways leading to tumour cell proliferation, invasion, and angiogenesis [12, 13]. In a related

approach, small synthetic molecules are developed to interact with the intracellular

phosphorylation sites of these receptor tyrosine kinases to disrupt downstream signaling [14].

Examples include imatinib, which blocks a receptor tyrosine kinase that drives chronic

myelogenous leukemia [15], and gefitinib, which binds mutated HER-1 common in lung and

colorectal cancers [16]. Targeted agents are designed to block cellular pathways that are mutated

in specific cancer cell types, but have a broad circulation and can also cause systemic toxicity by

inhibiting the same pathways in normal cells. While several molecular targeted agents have been

4

approved for clinical use, there have been few clear victories. First, targeted agents must be

carefully matched to an individual patient’s tumour genotype [17]. Even then, resistant tumour

subpopulations with minor mutations to the targeted molecule either already exist or rapidly

emerge [18-20]. Ongoing research aims to combine cytotoxic chemotherapy with molecular

agents targeted to multiple targets and signaling pathways, in an effort to outstrip the genetic

heterogeneity and evolutionary potential of tumour cells [21-23].

For any anticancer drug to be effective in vivo, it must not only be lethal to tumour cells,

but it must also be delivered to all tumour cells in a bioavailable form at a cytotoxic

concentration. Solid tumours pose specific challenges for drug delivery, and even with the

reversal of molecular and cellular mechanisms of drug resistance, or the use of monoclonal

antibodies and small molecule inhibitors designed to target specific receptors and signaling

pathways, new therapies will still need to overcome these hurdles to reach tumour cells.

1.1.3 Tumour microenvironment

Any localized mass of cancer cells can be classified as a solid tumour. Anticancer drugs that

show strong cytotoxicity to cancer cells in vitro have rarely been curative against solid tumours

in the clinic, due in part to acquired drug resistance, but also due to the inability of

chemotherapeutic agents to access all of the cells in the tumour interstitium at high enough

concentrations to induce cell death [24]. In normal tissue, the highly organized vascular network

Figure 1.1: Barriers to drug delivery in the solid tumour microenvironment.

5

that provides every cell with nutrients also enables the efficient delivery of anticancer drugs,

resulting in lethal toxicity to rapidly proliferating healthy tissue before achieving significant

antitumour effect. In tumours, drug penetration is limited by several confounding factors specific

to the solid tumour microenvironment (Figure 1.1). Blood vessels that develop during tumour

angiogenesis lack hierarchical organization, and are often leaky due to gaps in the endothelial

lining, basement membranes, and surrounding smooth muscle [25]. These structural and

hierarchical abnormalities lead to weak pressure differences between arteries and veins,

increasing resistance to blood flow. Periodic reductions in flow result in transient vessel

shutdown and hypoxia occurring on a time scale of 15 to 25 minutes [26]. Rapid tumour

proliferation outgrows the vasculature, forcing blood vessels apart to create chronically hypoxic

regions where cells are greater than 100 μm away from the nearest vessel [27], prohibiting the

efficient delivery of nutrients or drugs by diffusion. Compressed, non-functional lymphatic

vessels are unable to drain hypoxic regions of glycolysis products such as carbonic and lactic

acid, leading to increased fluid pressure and decreased pH [24, 28]. Interstitial fluid pressure

exceeding the microvascular pressure inhibits the convection of drugs into the tumour [29], and a

dense network of collagen fibers in the extracellular matrix impedes transport of large molecules

through the tumour interstitium [30]. This combination of tumour-specific factors, in addition to

the avidity with which anticancer drugs bind to their targets, limits drug distribution primarily to

the rapidly cycling cells of the perivascular space, leaving distal cells untreated and capable of

repopulating the tumour between treatment cycles [30].

One approach for selectively improving microregional drug distribution in tumours is

modification of the tumour microenvironment. Specific strategies include modification of

tumour blood flow, pH, oxygenation, interstitial fluid pressure, and vascular permeability by

either pharmacologic or physical means [24, 28, 31]. Enzymatic digestion of tumour

extracellular matrix components such as collagen and hyaluronan has been shown to facilitate

the diffusion of drugs into tumours as well as convection of macromolecules by reducing

interstitial fluid pressure [32, 33]. Anti-angiogenic agents such as VEGF inhibitors can be

combined with cytotoxic chemotherapy either to destroy endothelial cells of tumour vasculature

to deprive cancer cells of nutrients [34], or to prune immature and ineffective blood vessels while

sparing functional vessels to temporarily improve drug delivery and chemotherapeutic efficacy

6

[31, 35]; however, these approaches have been unsuccessful in the clinic [36]. An alternative

strategy is the development of hypoxia-specific agents to be used in combination with

radiotherapy or cytotoxic chemotherapy to target both well-oxygenated cycling cells and

dormant cells in hypoxic regions [37]. Physical modification of the tumour microenvironment

through mild heating has been shown to affect tumour blood flow, oxygenation, vascular

permeability, immunogenicity, and pH [38], as discussed in Section 1.2.

Another approach that has been developed for low molecular weight agents with short

half lives is to use macromolecular drug delivery systems to take advantage of the physiological

properties of solid tumours to increase antitumour effect while reducing systemic toxicity [39].

Nanoparticle drug carriers on the order of 50 to 400 nm conjugated to conventional anticancer

agents can slowly extravasate from gaps between endothelial cells into the tumour interstitium,

where their size and the tumour’s lack of functional lymphatics prevent them from being cleared

[40]. Drug carriers with long circulation times can thus preferentially accumulate in tumours

through what is known as the enhanced permeability and retention (EPR) effect [41], achieving

tumour concentrations up to 10-fold higher than normal tissue [41]. More than 20 such

passively-targeted nanoparticle therapeutics are now in clinical use [42]. However, their benefit

has typically been modified drug toxicity rather than improved antitumour effect, as their large

size inhibits drug transport in the tumour microenvironment [43].

Ongoing research aims to further improve the specificity of this passive tumour targeting

using drug carriers with targeting moieties such as antibodies or peptides [44]. By attaching

tumour-specific ligands, nanoparticles can be made to bind more avidly and specifically to their

target; however, the increased size and biological reactivity of these nanoparticles may reduce

their penetration in solid tumours [31]. For some chemotherapeutic agents, the increased

selectivity of passively or actively targeted nanoparticle drug delivery systems provides reduced

systemic toxicity allowing increased tumour dose [31]. It has been proposed that using

nanoparticle-mediated delivery to achieve high concentrations of anticancer drugs in tumours

could increase drug efficacy by overcoming some of the genetic and cellular mechanisms of drug

resistance to antineoplastic agents [45].

7

1.1.4 Doxorubicin

One of the most widely used chemotherapeutic agents that would benefit from improved drug

delivery is the anticancer drug doxorubicin. Doxorubicin, also known as Adriamycin, is an

anthracycline antibiotic derived from daunorubicin, which was isolated from the pigment-

producing bacteria Streptomyces peucitis [46]. Doxorubicin is active against leukemias and many

solid tumours, including breast, endometrial, ovarian and bladder cancers, as well as adult and

pediatric sarcomas, either alone or more commonly in combination with other drugs [47].

Typically administered as infusions of 50-75 mg/m2 repeated every three weeks, the dose and

frequency of administration are limited by acute myelosuppression, the occurrence of which is

correlated with the integral of plasma doxorubicin concentrations over time [48]. Repeated

doxorubicin administration is limited by chronic irreversible cardiomyopathy and congestive

heart failure that occurs with increasing frequency after a cumulative dose of 500 mg/m2 [49].

Doxorubicin cardiotoxicity is thought to be caused by the iron-mediated formation of free

radicals in myocardial mitochondria when exposed to peak plasma concentrations after injection

[50-52]. By reducing cardiotoxicity, continuous low-dose infusions and liposome formulations

that reduce peak plasma concentrations have been shown to increase the maximum tolerated

cumulative dose without affecting antitumour efficacy [51, 53]. It is believed that doxorubicin’s

antitumour effect is related to peak intracellular concentration [54] primarily through

topoisomerase II inhibition, but that several mechanisms may be involved [48]. At extracellular

concentrations of 1-2 μM that are expected for the first 10 minutes after intravenous bolus

injections, doxorubicin localizes in the nucleus of tumour cells where it binds to topoisomerase II

and prevents resealing of the double helix during DNA synthesis, leading to growth arrest and

apoptotic cell death. At extracellular concentrations above 2-4 μM (1-2 μg/ml), which exist only

within the first 5 minutes after bolus injection at the maximum tolerated dose, it is possible that

doxorubicin causes free radical formation and DNA cross-linking leading to cell death [48].

Intravenous bolus injections distribute rapidly into tissues with an initial half life of 5-10 minutes

[49], with peak plasma concentrations of 1-2 μM occurring immediately after injection, dropping

after one hour to concentrations of 0.025-0.25 μM (0.01-0.1 μg/ml), similar to those observed for

continuous infusion [48]. Doxorubicin is widely distributed in the body, with significant binding

8

to plasma proteins and tissue; full clearance after slow release from tissue-binding sites,

metabolism in the liver, and biliary tract excretion has a half-life of 24 to 48 hours [49].

Doxorubicin is fluorescent, with emission peaks at 554 and 585 nm when excited at

wavelengths of 470-490 nm (Figure 1.2 [55]). From fluorescence intensity, doxorubicin

concentrations can be quantified in extracts from homogenized tissue samples by fluorometry or

high-performance liquid chromatography (HPLC) with fluorescence detection [56, 57]. In

acidified ethanol buffer, this fluorescence increases linearly with doxorubicin concentration from

0.25-10000 ng/ml, with concentration quenching causing underestimation of concentrations

above 10 μg/ml [58].

Accurate quantification of the fluorescence of doxorubicin extracted from homogenized

tissue requires consideration of several modifiers of doxorubicin fluorescence including

homogenization technique, sample handling and storage, and tissue autofluorescence.

Doxorubicin can be extracted from homogenized tissue samples using organic solvents (0.3-0.7N

Figure 1.2: Fluorescence spectrum of doxorubicin. A) absorption spectra; B)fluorescence spectrum for Adriamycin (------) and its metabolite Adriamycinol (-- -- --).Purified samples were adjusted to optical density of 485nm = 0.6 in methanol forabsorption spectra. Fluorescence spectra of equimolar solutions in 95% ethanol, 0.6 N HClbuffer were obtained by excitation at 470 nm [55].

9

HCl in 48.5-70% ethanol [58-70], 2:1 v/v chloroform:isopropanol [64, 65, 71-79], 0.075N HCl

in 90% isopropanol [80-83]). However, its fluorescence decreases by drug decomposition into

non-fluorescent metabolites at a rate of 2% per hour at 25°C and 2% per day at 4°C [71],

remaining stable for prolonged periods at -20°C when stored in the dark [56, 57]. When loaded at

high concentrations in liposomes, doxorubicin fluorescence remains quenched until more than

95% of the drug is released [84, 85], with the fluorescence intensity of liposomes being

approximately 5% that of free drug [86]. HPLC shows that doxorubicin’s metabolites have

varying quantum fluorescence efficiencies [71, 87], but appear in small enough quantities that

their presence can be ignored [63, 88]. Converting fluorescence values to drug concentrations

requires calibration against a serial dilution of known quantities of doxorubicin, and correction

for autofluorescence in homogenized tissue samples [58].

Investigations of nanoparticle drug carriers for delivery of doxorubicin have focused on

increasing drug concentration in tumours or delaying tumour progression. However, the

relationship between these two endpoints is not direct [85]. One reason for a lack of correlation

between increased drug concentrations and antitumour effect is the microregional drug

distribution in solid tumours [30, 89]. As will be discussed in Section 1.3, liposome extravasation

is heterogeneous and the large size of liposomes prevents penetration away from tumour vessels

[31, 40]. Slow content release provides low extracellular concentrations of released drug [52],

and only small concentration gradients for diffusion away from tumour vessels [31]. For a

curative treatment, lethal cellular doxorubicin concentrations must be achieved in all tumour

cells, many of which may reside 100-200 μm from the nearest vessel [24]. In human non-small

cell lung cancer cells, Kerr et al [54] observed a correlation between cell death and intracellular

doxorubicin concentration (Figure 1.3), achieving 99% cell kill at approximately 60 μg/ml (110

μM), which required incubation for two hours at extracellular concentrations of 1-2 μg/ml (1.8-

3.7 μM). Intracellular concentrations increased faster and reached higher final concentrations

when exposed to higher extracellular concentrations, with the rate of cellular influx starting to

plateau at extracellular concentrations of approximately 2 μg/ml (3.7 μM) [54]. Even without

considering non-uniform drug concentrations in the tumour microenvironment, this is expected

to be achieved for only a few minutes after bolus injection of free doxorubicin [48]; much lower

extracellular levels, and thus longer incubation times, would be required for pegylated liposomal

10

doxorubicin. Putting these data in clinical context, Cummings and McArdle [73] measured

intratumoural concentrations of 819 ± 482 ng/g (0.819 μg/ml or 1.51 μM) in surgically excised

human breast cancer tumour samples at 30 minutes after intravenous injection of 25 mg/m2

doxorubicin. They observed lower concentrations in excised samples of gastric and colorectal

carcinoma, diseases which typically have a lower percentage of patients respond to single-agent

doxorubicin. Taken together, these findings suggest that antitumour effect can be improved by

increasing the extracellular concentration of doxorubicin around tumour cells, especially those

residing several cell layers from the nearest vessel.

Figure 1.3: Cellular uptake and cytotoxic effects of doxorubicin in human lung cancercells [54]. Left: Intracellular levels of Adriamycin after exposure to a range of externaldrug concentrations for up to 3 hours. Each point represents the mean of 4 experiments(±S.D.). Right: Relationship between clonogenic cell survival and intracellular drugconcentration. Each point is the mean of 4 experiments.

11

Figure 1.4: Doxorubicin distribution in tumour and normal tissue in MDA-MB-231tumour bearing mice at 10 minutes after injection. Photomicrographs of heart (A), kidney(B), liver (C) and tumour (D) tissues are shown. Doxorubicin (blue), blood vessels (red) andhypoxic regions (green). Fluorescence intensity gradients are shown in relation to distanceto the nearest blood vessel for all tissues (E) and on an expanded scale for tumours (F).(Scale bar represents 100μm) [95].

12

Fluorescence microscopy can be used to study gradients in drug concentration away from

tumour blood vessels either in vivo with window chamber tumour models [90, 91] or in

composite fluorescence microscopy images of tissue sections with drug fluorescence overlaid on

antibody-stained endothelial cells and hypoxic regions [92]. Examples of this type of quantitative

computerized microscopy have demonstrated in several mouse tumour lines [92-95] (Figure 1.4)

and in biopsies from human breast cancer patients [96] (Figure 1.5) that concentrations of free

doxorubicin decrease to half their perivascular concentration at 40-50 microns from the vessel,

unable to penetrate into hypoxic regions. In this thesis, similar techniques will be used to

investigate the ability to localize drug delivery with thermosensitive liposomes in a large animal

tumour model using focused ultrasound heating, and to evaluate the effect of this approach on

drug distribution in the tumour microenvironment.

Figure 1.5: Left: Microregional distribution of doxorubicin fluorescence in a cryosectionof a locally advanced breast cancer biopsy at 24 h after i.v. injection of doxorubicin (dose,90mg/m2 body surface). Lightest reds indicate highest doxorubicin concentrations. Middle:CD31 immunohistochemical staining combined with hematoxylin counterstaining of thesame biopsy area. Right: Mean nuclear doxorubicin fluorescence (averaged pixel intensityin arbitrary units) versus distance to the tumour boundary from the same biopsy.Doxorubicin fluorescence is concentrated near microvessels and the tumour islet boundary,with a steep doxorubicin gradient away from the tumour periphery [96].

13

1.2 Hyperthermia By increasing tissue temperature above 40°C, heat can be used to elicit a variety of cellular and

physiological effects that can either work in concert with other therapies or be used to induce a

direct cytotoxic effect. Two broad categories exist to classify these different therapeutic regimes.

Mild hyperthermia refers to tissue heating at temperatures of 40-45°C used primarily to enhance

and localize the effects of other therapies. Modest temperature elevations maintained for minutes

to hours can be used to increase blood flow, modulate immune response, induce chemo- or radio-

sensitization, increase vessel permeability, or trigger the action of temperature-sensitive drugs.

High temperature thermal therapy, sometimes called thermal ablation, is used to achieve

irreversible tissue destruction by applying temperatures greater than 50°C over timescales of

seconds to minutes.

1.2.1 Cytotoxic effects

The mechanism of cell killing by heat occurs through the inactivation of proteins caused by heat-

induced conformational changes [97, 98]. In the high temperature thermal ablation range (>

50°C), cellular and tissue structural proteins undergo rapid denaturation resulting in irreversible

structural and chemical changes, through the process of thermal coagulation [99]. In the mild

hyperthermia regime, protein unfolding temporarily renders cells vulnerable to other stresses

such as ionizing radiation and chemotherapy [100]. Cells respond to such damage by increasing

production of heat shock proteins, which act as molecular chaperones that bind to and protect

denatured proteins and guide their correct folding [98]. However, prolonged heating can

overcome this response leading to the formation of aggregates of denatured cellular proteins,

which can block essential cellular processes and trigger apoptotic cell death [101]. The cytotoxic

effects of hyperthermia are enhanced in tumours, whose disorganized vasculature is less

effective at removing heat; hypoxia and acidity in the tumour microenvironment increases

susceptibility to heat-induced cell death [102].

For mild hyperthermia, the rate of heat-induced cell kill is a nonlinear function of time

and temperature. At 43 to 45°C, heating must be maintained for more than an hour to achieve a

14

high degree of cell kill. Temperatures less than 43°C are ineffective at killing most cell types, but

are useful in the enhancement of other therapies. To compare heating regimens and predict

biological effect, the concept of a thermal isoeffective dose is commonly used to normalize

hyperthermia treatments to a number of minutes required at a reference temperature to achieve

an equal biological effect. The cumulative equivalent minutes at 43°C (CEM43) of a given

hyperthermia treatment is calculated as [103]:

( )∑ Δ=° Δ−

tT tR t43CCEM43 ,

⎩⎨⎧

°≥°<

=C43,50.0C43,25.0

TT

R [Eq. 1]

where tΔT is the average temperature during time tΔ and R is the number of minutes required to

compensate for a 1°C temperature change either above or below 43°C. R summarizes

experiments measuring the logarithmic time-temperature dependency of cell death in a variety of

human and animal tissues, which shows a breakpoint in the rate of cell kill at around 43°C.

Above 43°C, cell death approximately doubles for every 1°C temperature rise (R = 0.43-0.72),

and below this point it decreases by a factor of 4 (R = 0.125-0.25) [104]. Thermal dose

thresholds for various degrees of thermal damage have been measured in many tissues [104,

105]. The onset of thermal damage occurs in muscle at 30 CEM43 [106], while complete

necrosis occurs in various tissues at 50-240 equivalent minutes [104]. A threshold of 240

CEM43°C is commonly used as a conservative treatment goal to ensure thermal coagulation

[106, 107]. In clinical hyperthermia treatments, temperatures and thermal dose vary across the

tumour, making it useful to measure the thermal dose achieved by a certain fractional tumour

volume [102]. The thermal dose calculated from the time series of temperatures which 90% of

the target exceeds ( T90CCEM43° ) is a commonly used indicator of treatment effect, measuring

both the level of heat exposure and the coverage of the target region [108-110].

1.2.2 Chemosensitization

Hyperthermia enhances the cytotoxicity of anticancer drugs through a variety of mechanisms

[111-113]. Mild heating has been shown to inactivate proteins involved in repairing

chemotherapy-induced DNA damage [98]. Also, heat-induced inhibition of the transmembrane

15

drug efflux proteins responsible for multiple-drug resistance allows drugs to remain within the

cell at effective concentrations [100, 111]. For alkylating and other adduct-forming agents, the

increase in kinetic energy provided by hyperthermia accelerates reaction rates related to drug

action [112]. However, these effects vary widely between chemotherapeutic agents, and are

sensitive to heating duration, temperature, and drug dose, as well as the relative timing of

hyperthermia and drug administration [114]. For example, infusion of cisplatin or bleomycin

during 30 minutes of heating to 43°C shows thermal enhancement ratios of 2 to 4 over drug or

heat alone [115, 116]. However, for certain anticancer drugs including doxorubicin and

antinomycin D, if hyperthermia is administered before the drug, activated heat shock proteins

protect targeted cellular proteins, thereby decreasing therapeutic effect [113]. For several agents,

including doxorubicin, hyperthermia sensitization exhibits a temperature-threshold effect, with

strong thermal enhancement for thermal exposures of 30 min at 43°C vs. 41°C [111, 116].

Doxorubicin’s thermal chemosensitization is especially pronounced for local doxorubicin

concentrations higher than 10 mg/kg [111], which would be lethal if delivered systemically but

could be administered using localized drug release [45]. In recent clinical trials, hyperthermia

has improved treatment outcomes when combined with liposomal doxorubicin and paclitaxel for

patients with locally advanced breast cancer [117], and with combination etoposide, ifosfamide,

doxorubicin chemotherapy in the treatment of soft-tissue sarcoma [118].

1.2.3 Blood flow

In normal tissue, blood vessels respond to localized heating by the relaxation of vascular smooth

muscle cells, causing vasodilation and increased flow of normothermic blood to the heated

region [119]. This tissue-dependent thermoregulatory response can cause blood flow to reach 2-

20 times its resting rate [120], increasing with temperature up to a threshold at which vascular

damage occurs and blood flow is rapidly decreased [121]. Angiogenic tumour vessels have

abnormal smooth muscle cells and innervation, increasing their susceptibility to heat-induced

cell death and reducing their ability to autoregulate [121]. As a result, tumour vessels undergo

heat-induced decreases or small increases in tumour blood flow, and have a lower threshold for

vessel collapse. In a rabbit ear window preparation, blood flow in normal tissue increased with

heating up to 45.7°C for 60 minutes, while vessels in VX2 tumours experienced shutdown

16

following 30 minutes at 43°C [121]. In rat mammary carcinomas, 30 minutes of heating to 40.5-

43.5°C caused immediate increases in tumour blood flow and oxygenation that persisted for 24

hours, while 60 minutes at 42.5°C or higher caused acute vascular shutdown [122]. Comparing a

variety of other rodent tumours, heating to 40-42°C causes immediate increases in tumour blood

flow and oxygenation that persist for up to 24 hours, with acute vessel damage occurring at 43°C

or higher held for 30 minutes or more [119, 122]. Human tumour cells and tumour vasculature

demonstrate higher resilience to thermal damage when compared to rodent tumours [119, 123].

In human tumours, the ability of blood flow to cool tissue increases with temperature beyond 42-

43°C [124, 125], and even transient heat-induced decreases in blood flow occur only at

temperatures greater than 44°C [124]. During ultrasound hyperthermia, tumour blood flow was

found to either stay constant or increase at least transiently during 30-60 minutes of heating to

42.5°C, with increases starting approximately ten minutes after the onset of heating [126]. More

than most experimental tumours, the rabbit VX2 carcinoma is similar to human cancers as it

demonstrates high resting blood flow in its outer shell that increases as the tumour grows [127],

and the blood flow rate can transiently double when heated for 30 minutes at 43°C [121], making

it a good model to study the effects of hyperthermia on tumours [120] (Figure 1.6).

Figure 1.6: Effect of hyperthermia on blood flow in normal granulating tissue (solid) andVX2 tumours (dashed) in rabbit ear window preparation. Left: Peak flow rates recordedduring 60 minutes of heating at various temperatures. Right: For each temperature, thetime at which peak flow rate was reached. Blood flow increases with temperature up to athreshold at which vascular damage occurs and blood flow is dramatically reduced. Theeffect is stronger and the temperature threshold is higher in normal vs. tumour tissue [121].

17

1.2.4 Vessel Permeability

Angiogenic tumour vasculature is known to be ill formed and chaotic, with regions of higher

permeability than that of normal tissue, heterogeneously distributed along tumour microvessels

[40]. The “leakiness” of tumour vessels is exploited by nanoparticle drug carriers such as

liposomes, designed to preferentially leak out of tumour vessels and to achieve localized

accumulation in the tumour [128]. In animal tumours, hyperthermia has been shown to enhance

nanoparticle extravasation through changes in pore size [129, 130], where endothelial cells suffer

a transient disaggregation of microtubular proteins that acts to change the morphology of the

cytoskeleton causing the cell to “shrink”. This increases the occurrence and size of pores in the

vessel wall, enabling and increasing extravasation of normally impermeable particles as large as

400 nm [129, 131, 132]. For small particles that are inherently permeable to leaky tumour

vasculature, these increases in pore size have little effect on their rate of extravasation [133,

134]. For larger particles, exposures as short as 10 minutes can enhance permeability for several

hours [62, 84, 86, 135], which corresponds to the time required for heat shock protein-mediated

reassembly of microtubular proteins [130]. In rodent tumour studies, 30-60 minutes at 42°C has

been shown to increase extravasation up to 50 times [85], with both the maximum particle size

and rate of extravasation increasing with thermal dose up to the threshold for vascular collapse

[130]. By increasing the amount of nanoparticle extravasation from tumour vessels, mild

hyperthermia can selectively increase drug delivery to tumour cells.

1.3 Hyperthermia mediated drug delivery

1.3.1 Liposomes

One of the most effective drug carriers to take advantage of the enhanced permeability and

retention effect has been the liposome, a spherical vesicle enclosed by a lipid bilayer. While

liposomes can be made to carry hydrophobic drugs by insertion into the lipid bilayer, they are

more efficient carriers of hydrophilic drugs, which can be loaded into the aqueous core [136].

18

Initial liposome formulations of hydrophilic drugs failed to achieve tumour accumulation

due to rapid clearance from the bloodstream, with liposomes either forming rapidly eliminated

aggregates or being quickly marked for destruction and taken up in the spleen and liver by

phagocytic cells of the reticuloendothelial system [136]. This marking, or opsonization, occurs

through the adsorption of circulating plasma proteins to the lipid bilayer. Liposome elimination

can be delayed by incorporating long polymer chains of polyethylene glycol (PEG) into the lipid

bilayer [137]. Pegylation (4-5 mol%) creates a boundary that physically blocks liposome

aggregation [138] and binding of opsonization proteins [139, 140] to increase circulation times

from a few minutes to in some cases several days [141, 142].

In animal models, sterically-stabilized liposomal doxorubicin has shown the ability to

preferentially accumulate in tumours over 1-2 days [143, 144], demonstrating increased

accumulation in the perivascular tumour space, followed by prolonged drug release over a period

of 1-2 weeks as the drug overcomes transmembrane pH, electrostatic, and thermodynamic

gradients to diffuse out of the liposome, or the lipid bilayer is actively disrupted by lipoproteins

or macrophages [136]. Liposomes providing this altered biodistribution of doxorubicin

demonstrated increased antitumour activity with lower injected doxorubicin dose [145] and

decreased occurrence of dose-limiting cardiac toxicity [51, 144], but showed new dose-limiting

cutaneous toxicities due to liposome accumulation in the skin and extremities in what is known

as hand-foot syndrome [146]. For combined treatment of pegylated liposomal doxorubicin with

localized radiofrequency thermal ablation (70°C for 5 minutes), non-lethal heating at lesion

boundaries enhanced doxorubicin delivery at the margins, thereby increasing extent of

coagulation and lethal effects of doxorubicin [147-149].

Clinically, liposomal drug delivery offers a modified toxicity profile with decreased

cardiotoxicity, but little improvement in therapeutic effect [31, 43]. Pegylated liposomal

doxorubicin (Doxil, or Caelyx) [150] has demonstrated favorable toxicity profiles for AIDS-

related Kaposi’s sarcoma [151] as well as metastatic breast [152] and platinum-sensitive

epithelial ovarian cancer [153], but little improvement in antitumour efficacy over standard

chemotherapy regimens. Combination with hyperthermia shows promise for locally advanced

breast cancer [117], while a trial with whole abdomen hyperthermia for platinum-resistant

19

ovarian cancer showed a high occurrence of adverse events with no benefit over standard

treatment [154].

One reason for the limited improvement in antitumour effect is the sensitivity of

liposome extravasation to heterogeneities in tumour vascular permeability. While many studies

have demonstrated increased liposome accumulation in preclinical tumour models [143, 144],

liposome extravasation is heterogeneous within individual tumours due to large variations in

vascular permeability [155, 156], and widely variable between tumour types [128]. Furthermore,

while the size of liposomes enables their selective extravasation and sequestration in tumours, it

limits the penetration depth of liposomes to one or two cell layers from the nearest blood vessel

[156, 157]. In order to limit systemic release and retain drug until it accumulates in a targeted

tissue, liposome bilayers are designed to degrade more slowly than the accumulation half-life,

releasing drug over several days [136]. However, slow drug release results in low extracellular

concentrations and small concentration gradients; as a result, released drug has equal tendency to

diffuse back into the vasculature as it does out into the tumour [157]. In tumour regions with

limited vessel permeability, these factors lead to inadequate concentrations of bioavailable drug

and limited antitumour effect.

To increase tumour selectivity and antitumour efficacy, significant effort has gone into

pharmacologically targeting liposomes and other nanoparticle drug carriers to tumours by the

addition of monoclonal antibodies or other ligands such as DNA aptamers that bind specifically

to factors expressed by tumour cells [44], or to release their payload when exposed to factors

specific to the tumour microenvironment, such as pH, enzymes, or cancer-specific cell surface

antigens [158, 159]. These strategies improve the specificity of tumour specific accumulation,

thereby reducing toxicity to other liposome accumulation sites such as the liver, spleen, and skin

for a given systemic dose, but still rely on extravasation and nanoparticle transport [44]. Specific

ligand binding to tumour cells makes the treatment even more sensitive to variability in vascular

permeability, impeding extravasation and transport in the tumour extracellular matrix [31, 43].

Other methods aim to localize the release of encapsulated drug from liposomes or other

nanoparticle drug carriers by externally applied physical triggering mechanisms [158], such as

20

ultrasound pressure fields [160], light [161], magnetic fields [162], and heat [163]. Physical

release mechanisms are advantageous over methods depending on inherent tissue properties

(tumour-specific pH, enzymes, or cell surface antigens [158, 162]) in that they can be easily

modified to desired conditions. However, they are limited by their inability to target occult

microscopic metastases and systemic disease.

The use of local hyperthermia as a drug release mechanism is especially attractive

because of its ability to enhance the effectiveness of radio- and chemotherapy, for which it is

already used in the clinic [164]. Temperature-sensitive liposomes and other thermally-triggered

drug delivery systems with rapid drug release rates, including thermal-sensitive polymers [165,

166] and solid lipid nanoparticles [167], make it possible to release drug in a heated region as the

drug carriers circulate through the tumour vessels. Intravascular release can take advantage of

high vascular concentrations of liposomes continually entering the tumour, acting as a

continuous infusion delivering large amounts of free drug at the tumour site. Compared to

strategies where hyperthermia is used only to augment the passive accumulation of drug carriers

in the tumour, this direct mechanism retains the selectivity of liposomes, while increasing overall

tumour drug delivery and reducing sensitivity to variability in tumour vasculature.

1.3.2 Thermosensitive liposomes

By combining various proportions of dipalmitoyl phosphatidylcholine (DPPC) and other

phospholipids of different chain lengths and thus melting temperatures, temperature-sensitive

liposomes (TSL) can be designed with a liquid-crystalline transition temperature above

physiological temperature but attainable with mild hyperthermia (40-43°C) [168-170]. Upon

melting, these first-generation thermosensitive liposomes become leaky due to disorder at solid

and fluid domain boundaries (so-called “grain boundaries”), releasing 50% of encapsulated drug

over 30 minutes at 42°C [168]. In vivo, thermosensitive liposomes delivered large enhancements

in tumour drug deposition over free drug or non-thermosensitive liposomes [76, 84, 86, 135].

However, intravital microscopy studies demonstrated that tissue drug concentrations increased

linearly over 30-60 minutes of heating, with slow heat triggered release occurring only after

liposome extravasation [86]. The application of two hyperthermia doses, one before infusion to

21

increase blood flow and vessel permeability, and one several hours later to release drug from

liposomes accumulated in the tumour, could increase effect over a single dose [62]; however,

pre-heating may induce thermotolerance, reducing subsequent drug efficacy against tumour cells

[115]. With slow release rates, these enhancements in tumour drug concentration remained

highly dependent on hyperthermia-mediated liposome extravasation rather than triggered release

[76].

To reduce release time from minutes to seconds, thereby enabling intravascular release,

Needham et al developed lysolecithin-containing temperature-sensitive liposomes (LTSL) [171].

Traditional thermosensitive liposomes release drug slowly through grain boundaries that allow

leakage during the phase transition of bilayer lipids. So-called lyso-thermosensitive liposomes

containing approximately 10 mol% lysolipid (lipid with one fatty acid chain) in their bilayer take

advantage of the tendency of lysolipids to associate into micelles rather than liposomes [172]. As

the lipid bilayer is raised to its melting temperature, the lysolipid becomes free to diffuse

laterally within the bilayer, wrapping inward at grain boundary defects to form a micelle-like

lining. These lysolipid-lined nanopores are stabilized by pegylated lipids in the bilayer, further

enhancing permeability to particles up to 5 nm in diameter [172]. Recent formulations of this

type release 80% of their encapsulated drug in 20 seconds at 41.3°C [173]. The structure and

temperature sensitivity of this formulation are depicted in Figure 1.7 [172, 174-176].

In mice bearing implanted human squamous cell carcinoma xenografts (FaDu) treated

with water bath heating to 42°C and doxorubicin as either free drug, conventional liposomes,

passively thermosensitive liposomes, or LTSLs, induced curative effects in only the LTSL-

treated mice [76, 173]. LTSL-treated mice also had 20-30 times higher intratumour doxorubicin

concentration than free drug and 5 times more than conventional pegylated liposomes, with more

widespread drug deposition [76]. Enhancements of 6-15 times have since been reported across

additional rodent tumour models using a variety of heating techniques [68, 177-180].

Pharmacokinetic and biodistribution data from tumour-bearing dogs suggest that release from

lyso-thermosensitive liposomes achieves increased drug concentrations in tumours vs. non-

thermosensitive liposomes, with a similar toxicity profile to free drug [181].

22

Figure 1.7: Structure and temperature-sensitivity of liposomal drug carriers. A)Doxorubicin release from lysolipid-containing temperature sensitive liposomes (LTSL).Upon pH-gradient loading in liposomes at high concentrations, doxorubicin stacks intofibers that condense into bundles [174]. Above the lipid melting temperature, doxorubicinis rapidly released through lysolipid-stabilized pores. B) Cryogenic transmission electronmicroscopy of doxorubicin fiber bundles in LTSL before heating [175]. Scale bar 200 nm.C) Release of doxorubicin from LTSL vs. time and temperature. Release rate is highest atthe critical temperature of the liposomes [176]. D, E) Mechanism for enhanced release ratefrom LTSL. In traditional DPPC temperature sensitive liposomes (D), triggered releaseoccurs at the melting temperature of DPPC (41.5°C) through interfaces between regions ofsolid and liquid lipids in the liposome bilayer. In temperature sensitive liposomescontaining 10% MSPC lysolipid (E), heating to the melting temperature allows MSPC toform micelle-like structures at DPPC grain boundaries, acting as pores for doxorubicinrelease. Pores are thought to be stabilized by adding 4-5% DSPE-PEG2000 to the bilayer,which also enhances liposome circulation time [172].

23

It has been hypothesized that the increased tumour drug concentrations and antitumour

effect of lyso-thermosensitive liposomes may be caused by rapidly-triggered intravascular and

perivascular release causing significant antivascular effects [79, 182, 183] and high local

concentrations that drive free drug into the tumour interstitium by diffusion [76]. Figure 1.8

illustrates this hypothesis. This intravascular release has the potential to overcome the obstacles

of vascular heterogeneity and limited penetration associated with liposome extravasation, while

traditional thermosensitive liposomes formulations with slower release rates depend on drug

release after accumulation in the perivascular interstitial space. This hypothesis of intravascular

drug release from LTSL is supported by data obtained using window chamber tumour models

demonstrating release of a fluorescent dye from liposomes within the vasculature, and

subsequent extravasation into the tumour interstitium [184]. If the same mechanism occurs for

doxorubicin in solid tumours, it is possible that lyso-thermosensitive liposomes can deliver high

intravascular drug concentrations capable of driving cellular drug uptake and increasing drug

penetration further from vessels, increasing drug exposure of all tumour cells.

Fast, intravascular release necessitates the use of precisely controlled hyperthermia to

achieve localized, tumour-specific drug delivery. Inadequate heating results in insufficient drug

release, while excess heating can result in vascular damage and shutdown, preventing liposomes

from reaching the target region. MRI-guided focused ultrasound is a promising hyperthermia

technique for localized drug delivery because it provides precisely controlled energy deposition

and accurate dosimetry of achieved temperature elevations for selective, noninvasive heating of

deep-seated tumours.

1.3.3 Clinical Experience

Based on promising preclinical results, Needham’s temperature-sensitive liposome technology

has been licensed and brought to clinical trials by the Celsion Corporation. In an ongoing phase

I/II trial for breast cancer recurrences at the chest wall [185], LTLD is administered

intravenously prior to mild heating with externally-applied microwave devices, which achieve

temperatures of 40 to 42°C in diffuse superficial tumours [186]. After the 9-patient phase I

portion determined a maximum tolerated dose of 50 mg/m2 administered as six combined LTLD

24

and hyperthermia treatments at 21-day intervals, the 100-patient phase II portion aims to

determine the rate of durable (longer than 3 month) complete local response.

Further clinical experience is available for liver tumours, where invasive radiofrequency

(RF) thermal ablation destroys the central tumour mass with temperatures greater than 60°C,

while peripheral heating below the thermal coagulation threshold releases doxorubicin from

Figure 1.8: Localized intravascular drug release from thermosensitive liposomes.

25

circulating liposomes [187]. Drug release at the boundaries of the thermal lesion increases the

size of treatable tumours and eliminates microscopic deposits of residual tumour cells at tumour

margins. In a Phase I trial involving 24 subjects who received a 30-minute infusion of LTLD 15

minutes before radiofrequency ablation of hepatocellular carcinoma or metastatic liver cancer,

patients receiving at least the maximum tolerated dose of 50 mg/m2 had an increased median

time to treatment failure over those receiving lower doses of LTLD (374 vs. 80 days) [187, 188].

Thermodox was demonstrated to have a circulation half-life of 18 h, longer than that of free

doxorubicin (5 min), but shorter than pegylated liposomal doxorubicin (Doxil, 55 h) [189].

LTLD demonstrated a favorable toxicity profile over other doxorubicin formulations, exhibiting

no signs of free doxorubicin’s dose-limiting cardiac toxicity or liposomal doxorubicin’s hand-

foot syndrome [187]. With encouraging Phase I results suggesting that the combination of LTLD

and RF ablation is safe and likely better than RF alone, Celsion was allowed by the FDA to

proceed directly into a multicenter Phase III trial [190]. The Phase III trial is expected to

determine whether progression-free survival is improved for RF ablation with LTLD vs. RF

ablation alone among 600 patients with unresectable (larger than 3.0 cm) hepatocellular

carcinoma.

Preliminary evidence from clinical trials suggests that combining thermosensitive

liposomal doxorubicin with localized heating is safe, has a favorable systemic toxicity profile

relative to free or liposomal doxorubicin, and might enhance local control over heat alone.

However, the localized delivery of high doses of doxorubicin requires accurate localization of

heating. The following section identifies several limitations with current clinical hyperthermia

techniques, and proposes the use of MRI-controlled focused ultrasound as a means of

noninvasively delivering precise, localized hyperthermia for thermally-triggered drug delivery.

1.4 MRI-Controlled Focused Ultrasound Hyperthermia

1.4.1 Clinical Hyperthermia and Thermal Therapy Techniques

In the clinical use of mild temperature elevations to enhance and localize the effects of other

therapies, it would be desirable to have a noninvasive technique capable of accurately controlling

26

spatial energy deposition in tissue. The lack of devices capable of reliably heating tumours to

prescribed temperatures has limited the range of applications for hyperthermia in the clinic [191].

Thermal therapy has instead been used primarily in the context of localized percutaneous tumour

ablation with interstitial radiofrequency, microwave, or laser applicators [192]. RF applicators

operating at around 470 kHz deposit energy through resistive loss near the electrode as current is

passed from the applicator tip through tissue to a ground return located either on a second needle

within the target or on an external grounding pad [193]. Energy absorption drops rapidly with

distance from the applicator tip, with heating by thermal conduction at the lesion periphery to

generate thermal lesions of up to 3 or 4 cm in diameter using multi-tined electrodes [194]. Tissue

desiccation at temperatures approaching the vaporization threshold decrease tissue conductivity

at the applicator tip, limiting energy penetration in tissue; active device cooling and direct saline

injections have been investigated to overcome these effects [192]. Interstitial microwave

antennae operating at 915 or 2450 MHz are less sensitive than RF to variations in the thermal

and electrical conductivity of tissue and the presence of blood vessels, using radiative

propagation of electromagnetic waves and dielectric energy absorption to produce larger ablation

zones with a broader region of active heating [195]. Laser fibers provide ablation zones of 1 to 2

cm in diameter with inherently MR-compatible devices that have found use in the treatment of

small tumours [196, 197], most notably in the brain [198, 199]. However, at optimal wavelengths

of 800-1100 nm light has a penetration depth of only 7.5 mm, and early tissue coagulation and

charring near the device causes further reductions in penetration depth [200].

To achieve regional or systemic effects, uniform mild heating of large areas can be

delivered to the body using thermal conduction by contact with heated fluids, air, or blankets, or

by infusion of a heated fluid into the blood to achieve systemic temperatures of up to 42°C [201].

Microwave electromagnetic radiation at commonly used frequencies of 433, 915, or 2450 MHz

can be applied with external waveguides or antenna applicators, and is well-suited for

noninvasive hyperthermia of broad superficial regions at 40-42°C to a depth of not more than 3

cm [202, 203]. Regional hyperthermia of deep-seated tumours can be achieved noninvasively

using antenna arrays operating at more penetrating frequencies between 100 and 150 MHz [201].

At these frequencies, electronically phased applicators have some ability to focus

electromagnetic energy deep into the body, but with wavelengths of 20-30 cm and heterogeneous

27

energy propagation dependent on patient anatomy, it is said that one is forced to “dump” energy

in a large region “and pray” that the tumour will experience selective heating due to its low

perfusion relative to nearby critical organs [204, 205]. Recently, hybrid electromagnetic

hyperthermia-MRI systems have been developed to use MR thermometry to monitor

temperatures achieved during heating [206, 207]. Spatially resolved temperature information

enables real-time adjustment of electromagnetic focusing to improve tumour coverage and

reduce heating in unintended hotspots, within the physical limitations of this modality.

Ultrasound can be used to heat tissue, inducing particle vibrations whose energy is

partially absorbed as the wave propagates through the body. Due to its short wavelength at

frequencies that penetrate well in tissue, ultrasound can be focused to noninvasively produce

therapeutic temperature elevations in a millimeter-sized region several centimeters beneath the

skin [208-210]. Ultrasound can be applied with either interstitial or external devices, and used for

either mild heating or thermal ablation. Here we focus on noninvasive externally-focused

ultrasound for mild hyperthermia.

1.4.2 Focused Ultrasound

Ultrasound is a mechanical wave that propagates within a medium inducing particle vibrations at

a frequency above 20 kHz which are capable of eliciting a variety of biological effects. These

include the generation of heat through absorption, mechanical displacements due to radiation

pressure, and cavitation, the creation and manipulation of gas bubbles in a medium. As

ultrasound travels through tissue it is attenuated mostly through absorption (>80%), as the

energy of the particle vibrations are converted into heat. By geometrically focusing the waves, a

focal point on the order of the ultrasound wavelength (approximately 1 mm) can be obtained

with high intensity and energy deposition, capable of heating deep-lying tissue with negligible

temperature increase between the transducer and its focus. The millimeter-sized focal volume is

attractive for localized tissue ablation but presents challenges for hyperthermia, where volumes

of several centimeters might need to be heated to uniform temperatures. One method proposed to

heat regions larger than the size of the ultrasound focus is to move the focal point along a

circular trajectory, taking advantage of the thermal conductivity of tissue to heat the tumour

28

interior [208, 211, 212] (Figure 1.9). The ultrasound focus can be scanned during continuous

sonication to heat large tumour volumes by either mechanically repositioning a single element

transducer, or by electronically steering the focus generated by the hundreds of small,

independently-driven transducer elements in a phased array [213].

The change in temperature resulting from a given energy deposition can be

approximately modeled using the Pennes bioheat transfer equation (BHTE) [214]:

( )

cp

TTwCTtTC ABT ρ

ακρ2

2 +−−∇=∂∂ [Eq. 2]

Figure 1.9: Strategies for using thermal conduction to heat large tissue volumes usingfocused ultrasound. Top: Moving a small focal point along the periphery of a targetedregion allows thermal conductivity of tissue to heat the target interior. Bottom: Withexternally focused ultrasound transducers, the focus can translated during continuoussonication by either mechanically repositioning a single element transducer (left), or byelectronically steering the focus generated by the hundreds of small, independently-driventransducer elements in a phased array (right).

29

Where the change in temperature T∂ depends on the tissue density ρ , specific heat TC , and

ultrasound absorption coefficient α , as well as the tissue speed of sound c and applied acoustic

pressure amplitude squared 2p , subject to heat diffusion defined by the tissue’s thermal

conductivity κ , and heat convection through the blood with perfusion rate w and specific heat

BC . Spatial, temporal, and temperature-dependent variations in tissue absorption, diffusion, and

perfusion create difficulty in making accurate predictions of temperature elevation for a given

applied power [208, 215, 216]. During short bursts of energy, the temperature rise is mostly

perfusion independent [217], but when the heating duration is long, perfusion significantly

reduces the achieved temperature, and is known to vary with duration of applied heating [126,

216]. This difficulty in predicting temperature elevation can only be overcome by regulating the

applied energy using feedback from temperature measurements made during heating [212, 218-

220].

While the ability to focus acoustic energy into tissue has been understood for more than

fifty years [221-223], this noninvasive technique struggled to find clinical use due to a lack of

image guidance. Complex tissue interfaces affect the position of the ultrasound focus, and

undesired hot spots can occur if there is highly absorbing bone or highly reflecting air-filled lung

or bowels in the beam path. These uncertainties necessitate visualization of target regions,

guided localization of energy delivery, monitoring of immediate local effects, and evaluation of

treatment outcome. Magnetic resonance imaging (MRI) is well-suited for each of these tasks, as

it is non-ionizing, has excellent anatomic resolution using a variety of contrast mechanisms, and

has the capability to noninvasively measure tissue temperature. MR thermometry enables precise

localization of the ultrasound focus by visualizing the 3-5°C temperature rise achieved during

short, low-power sonications. For high-temperature thermal ablation, temperature maps acquired

during therapy can be used to calculate the deposited thermal dose and predict the extent of

thermal lesions. These rapidly acquired temperature measurements can also be used with

feedback control algorithms to automatically modulate power deposition for control of the spatial

temperature distribution over time. Clinically, MRI-guided focused ultrasound thermal therapy

[224-228] has been used to coagulate benign and malignant breast tumours [229-233], uterine

fibroids [234-236], liver cancer [237], bone metastases [238, 239], and brain tumours [240], as

30

well as deep brain structures responsible for specific neurological conditions such as chronic

pain [241], essential tremor [242], and Parkinson’s disease [243]. Lower temperature

applications (temperature elevations of 4-7°C) in preclinical development include spatial and

temporal control of transgene expression [244, 245], and as described in this thesis, locally

triggered release of contrast agents and anticancer drugs from thermosensitive drug carriers [68,

69, 163, 246, 247].

1.4.3 MR Image Guidance and Thermometry

1.4.3.1 Magnetic Resonance Imaging

Magnetic resonance imaging uses a combination of static and time-varying magnetic fields to

measure tissue anatomy and function through their effects on the magnetization of water protons

in the body [248]. In the presence of a strong external magnetic field, protons, or spins, tend to

align with the external magnetic field, B0. If the magnetization is out of alignment with B0, it

spirals along a helical pathway towards alignment, precessing about the axis of the magnetic

field at a frequency determined by its strength:

00 Bγω = [Eq. 3]

For water protons, the gyromagnetic ratio γ is 42.58 MHz/T, and for commonly used magnetic

field strengths B0 of 1.5 to 7.0 T, the resonance, or Larmor, frequency 0ω is in the

radiofrequency range (64 to 300 MHz).

Spins can be excited out of thermal equilibrium by the pulsed application of a time-

varying magnetic field B1. B1 is designed to oscillate in a plane perpendicular to the main field B0

and contains a narrow range of frequencies centered on 0ω , spiraling the precessing

magnetization away from the B0 axis into what is known as the transverse plane. This precessing

magnetization is the source of the NMR signal, producing an oscillating magnetic field that

induces a current in a coil placed near the sample. Spins in various tissues relax from excitation

back to thermal equilibrium at different rates through interactions with the local magnetic fields

31

of neighbouring protons and other molecules, giving rise to tissue-specific variations in signal

strength.

To localize the measured signal and create an image of the magnetization, an additional

set of coils is used to create linear spatial variations in B0. By applying such a gradient in one

direction and then restricting the range of frequency components in the B1 excitation pulse to a

subset of the resulting proton resonance frequencies, excitation can be limited to a tissue slice of

desired thickness. Within this slice, gradients in resonance frequency applied for a given duration

create a spatially varying phase, or in other words, a spatial frequency. For objects of limited

size, we can sample the voltage induced in the receiver coil after applying gradients in two

directions for a variety of durations, measuring signal over the entire range of spatial frequencies

required to fully describe the object in the spatial frequency domain, referred to as k-space. An

object domain image of the magnetization can then be recovered using an inverse Fourier

transform of the acquired k-space data.

Relaxation is highly tissue dependent. The time constant of an exponential function

approximating the rate at which the magnetization returns to alignment with B0 is called T1, and

the time constant describing the rate at which the transverse magnetization returns to zero is

called T2. The sequencing of radiofrequency excitation pulses, magnetic field gradients, and data

acquisition can be used to weight the measured signal by a desired contrast mechanism, allowing

many different tissue properties to be spatially mapped, including the density of protons in a

sample, their T1 and T2 relaxation rates, water diffusion, blood oxygenation, magnetic

susceptibility, and temperature. Several contrast mechanisms are useful in identifying cancer and

defining target boundaries for therapeutic interventions; in thermal therapy, the high spatial

resolution and soft tissue contrast of MRI can also be used to evaluate treatment response [249].

One approach is to observe the dynamics of injected contrast agents to identify regions of

decreased perfusion caused by changes in the vasculature after thermal therapy. With dynamic

contrast enhanced MRI using gadolinium-based contrast agents, thermal lesions appear

immediately after ablation as dark, non-enhancing regions on T1-weighted images [106] that

correlate well with the histological extent of thermal damage [250]. Another approach is to

32

directly identify coagulation-induced changes in tissue properties by changes in signal intensity

on T1, T2, or diffusion-weighted images [251, 252]. In particular, T2-weighted images depict

thermal lesions as regions of low signal intensity surrounded by a ring whose intensity increases

over the course of 1-2 hours after heating [106, 253] that correlates with histologically-identified

oedema and inflammation [254, 255]. MR is also the only clinically accepted method to

noninvasively quantify temperature changes deep within tissue. MR thermometry can be used to

identify the location of the ultrasound focus during low-power exposures [256], as well as to

monitor temperatures in the target and normal tissues during treatment [257]. This temporal and

spatial temperature information allows us to noninvasively estimate and control the delivered

time-temperature history or thermal dose, and is the most important feature of MRI for guiding

thermal therapy.

1.4.3.3 MRI Thermometry

Early hyperthermia systems used temperature feedback from arrays of interstitial thermocouple

probes [126, 220, 258, 259]. By adding an invasive aspect to an otherwise noninvasive

procedure, thermocouples and fiber-optic probes introduce the risk of toxicity or treatment

complication, while providing only a coarse sampling of the spatially varying temperature

distribution. Furthermore, thermocouples interact with the acoustic field to create temperature

measurement artifacts and distort the energy absorption pattern in tissue [260]. The development

of magnetic resonance temperature imaging has provided a noninvasive method to make

spatially-resolved temperature measurements for online guidance of hyperthermia [261].

Several MRI parameters can be used to measure temperature, including spin-lattice

relaxation time T1 [262], water diffusion [263], equilibrium magnetization [264], and the water

proton resonance frequency shift [265] which can be measured using MR spectroscopy [266] or

phase mapping [267]. For thermal therapy, the phase-difference proton resonance frequency shift

technique has several advantages over the other techniques and is by far the most commonly

used.

The physical basis for the proton resonance frequency shift is the subtle change in local

magnetic field around water protons caused by the decreasing number of hydrogen bonds per

33

water molecule with increasing temperature, due to the increasing mobility of water molecules

[265]. As temperature increases and the kinetic energy of water molecules overcomes the

strength of the hydrogen bonds between them, water protons dissociate from the oxygen atoms

of their neighbours and experience a greater shielding by their own electrons, thus experiencing a

smaller local magnetic field and hence Larmor frequency ω0. During the echo time between MR

signal excitation and measurement, the small frequency change results in an accumulated phase

offset proportional to the change in temperature. Images of these phase offsets can be acquired

from gradient echo MR image acquisitions. However, phase variations due to heating are much

smaller than the background phase variations that arise from the static magnetic field distortions

caused by the susceptibility of objects placed in the bore. To calculate temperature maps using

the PRF technique, phase differences arising from relative temperature changes can be isolated

from static phase variations by subtracting phase images acquired before and during heating

[267]:

oBTET

⋅⋅⋅−

=Δαγφφ 12 [Eq. 4]

This relationship between temperature change ( TΔ ) and phase difference ( 12 φφ − ) depends only

on the main magnetic field strength ( oB , often 1.5 or 3.0 Tesla), the prescribed echo time, (TE ,

typically 5-30 milliseconds) the gyromagnetic constant (πγ2

= 42.58 MHz/Tesla), and the proton

resonance frequency shift coefficient (α = 0.01 ppm/°C). The PRF shift coefficient is constant

for a wide range of temperatures (0 to 100°C) and is insensitive to tissue type for water-based

tissues, even after thermal coagulation [268]. Phase images can be rapidly acquired (typically 0.1

to 5 seconds) in multiple slices through the use of gradient echo pulse sequences, maintaining a

sufficient signal-to-noise ratio for typical temperature measurement uncertainty of less than 1°C

at 1.5 T. An example of temperature measurement using the PRF technique is shown in Figure

1.10. In this experiment, the temperature distribution transverse to an externally focused

ultrasound transducer was measured during sonication of a gel phantom, depicting the shape of

the heating pattern.

34

Figure 1.10: Temperature measurements made with MRI using the proton resonancefrequency shift phase-difference technique. A) Magnitude reconstruction of gradient echoacquisition depicts transverse slice through an externally focused ultrasound transducersituated in a bath of degassed, de-ionized water, as it heats a focal region in a gel phantomabove. B, C) Phase image reconstructions acquired before and after application ofultrasound energy. D) Subtraction of phase images in B and C, scaled to relativetemperature, identifying in this case a 16°C temperature rise caused by focused ultrasoundheating.

35

The biggest challenge with the phase subtraction technique is the time-varying change in

magnetic field pattern due to breathing and motion. Object motion in the image plane causes

misregistration of the image subtraction, while motion anywhere near the image plane can alter

the susceptibility distribution to produce artifactual phase changes that mask the temperature

signal. This poses major difficulties for temperature measurement near moving organs such as

the heart and liver, or in the vicinity of the inflating and deflating lungs. The phase difference

technique is also sensitive to slow variations in the main magnetic field; this magnetic field

“drift” can be corrected by subtracting the phase measured in image regions where the

temperature is known to be constant [268-271]. Another important limitation is that this mode of

temperature-dependence applies only to water protons and works well only in tissues with high

water content. In tissues containing a high proportion of lipid, lipid suppression can isolate

temperature-sensitive water signal, but the temperature-dependence of lipid susceptibility can

still modify the local magnetic field, confounding the PRF measurement [272, 273]. Finally,

because the technique measures a relative temperature change, it requires accurate temperature

information when the background images are acquired. Despite these limitations, phase

difference MR thermometry provides the ability to measure dynamic temperature distributions

noninvasively during thermal therapy, enabling accurate dosimetry and damage predictions. This

has been essential to the clinical translation of noninvasive ultrasound thermal therapy, and has

been used successfully in humans to guide thermal therapy in many tissues including the uterus

[235], brain [198, 240, 241], breast [229, 232], rectum [274], and prostate gland [275].

1.4.5 MRI-Controlled Focused Ultrasound

1.4.5.1 Motivation for feedback control

In a typical MR-guided treatment, anatomical MR images are used in treatment planning to

prescribe a target region to be heated to a desired temperature or thermal dose. Energy is

deposited with a predefined power, spatial deposition pattern, and duration, and the resulting

temperature elevation is measured with MRI, verifying treatment outcome by thermal dose or

other indicators of thermal damage. However, even when energy can be selectively delivered to

targets deep within tissue, spatial variations in energy absorption and blood flow make it

36

technically challenging to maintain temperature uniformity in a heated target over time.

Clinically, temperature elevations achieved with focused ultrasound for a given applied power

have been shown to vary by 30-40% both between different patients and between sonications for

one patient due to beam aberration at tissue interfaces and variations in energy absorption [235].

Temporally, temperature fluctuations with a standard deviation of 1°C result in a 25% increase in

the estimate of thermal dose [276], making it important to minimize temperature measurement

uncertainty [276-278] and actual temperature fluctuations [219]. Spatial temperature variations

are especially difficult to avoid in the presence of large blood vessels [279], and even more so

when heating with a small scanned focus, making spatial control of applied energy a necessity

[218, 219]. For thermally-triggered drug release from temperature sensitive liposomes, tissue-

specific heating is necessary to achieve locally enhanced drug delivery, as well as to avoid

systemic drug release and normal tissue toxicity [85]. Furthermore, precise control of

temperature elevation is required for optimal drug release without prematurely preventing drug

supply by damaging tumour vessels.

In an MRI-controlled approach, temperature measurements are made rapidly and used as

input to a feedback control algorithm that automatically updates the location, duration, and

power of energy deposition during treatment in order to track the prescribed target temperatures,

ensuring that a desired spatial and temporal heating pattern is achieved and maintained.

The temporal, spatial, and temperature resolution requirements for thermometry used in

closed loop feedback control depend on the type of treatment, and tradeoffs can be made. In

point-by-point thermal ablation, large temperature increases in small focal regions over short

durations require high spatial and temporal resolution, but the requirement for temperature

resolution is relaxed, as the primary goal is to exceed a thermal dose or temperature threshold.

For volumetric ablation where a large region is rapidly heated to high temperatures all at once,

temporal resolution is important, but the spatial resolution requirement is relaxed. Clinically, a

simple on-off control approach has been used to achieve uniform coagulation of uterine fibroids

using externally focused ultrasound [255, 280]. By electronically steering a focal point along

concentric circles and moving outwards from one circle to the next once the boundary

temperature reaches 57°C, this simple approach achieves precise, reproducible lesions up to 16

37

mm in diameter [281]. A similar threshold-based approach has been used for transurethral

ultrasound ablation of the prostate, where a planar ultrasound transducer heats the prostate from

within the urethra and is slowly rotated as the target boundary reaches a threshold temperature

[282]. However, for mild heating to enhance chemotherapy, small temperature elevations must

be maintained over a long duration, with conduction effects spreading the heat over time. In this

case, the requirements for temporal and spatial resolution are relaxed, but precise temperature

resolution is required, as fluctuations of 1-2°C can cause dramatic changes in biological effect.

For mild hyperthermia using focused ultrasound under MRI control, several closed-loop

feedback control algorithms have been studied in animals, while closed-loop thermal ablation

has been demonstrated in humans. In general, automated feedback control provides more

accurate temperature regulation than manual operator control, especially for systems with

temporally and spatially varying energy deposition patterns.

1.4.5.2 Scanned focused ultrasound hyperthermia

Prior to the introduction of noninvasive temperature measurement using MR thermometry, a

mechanically scanned focused ultrasound system was designed and used to treat over 175

patients [126], controlled by temperature feedback from a sequentially sampled array of

interstitial thermocouple probes [220, 258, 259, 283]. The system consisted of a modified

commercial diagnostic ultrasound imaging system mounted with four focused transducers (7-13

cm diameter, 25-35 cm radius of curvature, 0.5-1.0 MHz) whose beams overlapped at a single

acoustic focus. The entire transducer gantry could be rotated, tilted, and translated in three

dimensions by five stepper motors under computer control. To achieve uniform heating of a

region larger than the acoustic focus, the gantry and focus were scanned in single or multiple

concentric octagonal trajectories in a plane parallel to the focal plane at speeds of up to 50 mm/s

[220]. Parametric studies of mechanically scanned focused ultrasound hyperthermia offer several

recommendations for achieving uniform heating in large regions:

Peripheral energy distributions produce uniform temperature elevations with steep outer

gradients, and spatial intensity variation under feedback control can be used to overcome

heterogeneous tissue properties and perfusion [205, 208, 211, 212, 216, 259]

38

Angular temperature variations and periodic focal heating/cooling along the scan path are

minimized with fast scan times and low frequency transducers [211, 219]: for a 2 cm scan

diameter with a 1 MHz transducer, fluctuations of < 1°C can be achieved with 5 second scans at

perfusion rates of up to 20 kg/m3/s. 1 second scans typically eliminate changes in thermal dose

[212].

High perfusion rates decrease temperature magnitude and uniformity, but can be

overcome using fast scan times and concentric circular scans with focal diameter spacing [212,

216, 284, 285].

Continuous scanning causes pre-focal heating that increases with frequency and f-number

(ratio of radius of curvature to aperture size), and far-field heating that decreases at higher

frequency. Pre-focal heating increases with perfusion for shallow focal depths but skin cooling

can help. To reduce prefocal heating, increased acoustic windows can be achieved by tilting the

transducer for a similar effect as using a lower f-number [212, 284, 286].

Temperature maxima occur at muscle-bone interfaces as far as 5 cm outside the focus if

any part of the acoustic field is incident on bone during the scan. Far field maxima are reduced at

higher frequencies (i.e. 3 MHz) [215].

Undesired hot spots can occur at skin-air interfaces beyond the focus due to reflection of

the exiting beam and strong absorption in the skin, causing temperature elevations as high as 4-6

times those observed at the focus [287]. Ultrasound field intensity at skin-air interfaces can be

reduced using higher frequency and sharper focusing for faster attenuation and divergence

beyond the focus. Higher absorption at the focus relative to the skin can be achieved with non-

linear ultrasound propagation using short, high-power pulses. Reflection at the skin can be

reduced by using ultrasound gel to couple the exit beam out of the tissue into a water bag with a

rubber absorber on the other side. These issues are especially important for targets in the head

and neck, extremities, and in small animals.

With multi-element phased array transducers, uniform heating of 1-3 cm regions can be

achieved by rapid electronic switching (50 ms) between a series of multi-focal patterns [213] or

39

scanning of a single focus [288]. Fast, online control of focal spot size and location with no

moving parts makes phased arrays desirable for routine clinical applications, but complex,

expensive driving electronics and fixed geometry limit their use.

1.4.5.3 Proportional-integral-derivative control of focused ultrasound thermal

therapy

Control algorithms based on proportional-integral-derivative (PID) closed-loop feedback were

developed for the Arizona scanned focused ultrasound system with temperature measurements

from invasive thermocouples [259], and later generalized to use MRI thermometry [289, 290].

With discrete-time PID control, power applied to the heated region is adjusted every time a

temperature measurement is received, based on the difference between the predefined

temperature goal gT and temperature samples iT at times ni ..1= : [259]

( ) ( ) ( )nnD

n

iigIngP TTKTTKTTKP −+−+−= −

=∑ 1

1

[Eq. 5]

The proportional term (scaled by proportional gain, PK ) compensates for the current error

between the measured and target temperatures ( ng TT − ) to give faster response to disturbances,

with excessive values causing oscillation and instability. Integral gain IK reacts to the sum of all

previous errors, gradually adjusting the applied power to reduce steady-state error at the cost of

increased overshoot and settling time. Derivative gain DK acts to increase power when the target

temperature is changing, but differentiation of temperature measurements acts to amplify noise,

possibly leading to instability. Spatial temperature control can be achieved by using multiple PID

controllers to independently modulate the applied power or sonication time for each point along

a defined trajectory [258, 291]. Lin et al [259] measured spatial and temporal temperature

variations of less than 2°C in 15 mm diameter regions using control parameters based on

simulation results, with feedback from invasive temperature probes placed along the scan path.

Subsequent control strategies have taken advantage of thermal modeling and the spatial

temperature information provided by MRI thermometry to become more robust to spatial and

40

temporal variations in blood flow and energy absorption. In adaptive control algorithms, gain

terms are added or adjusted to replace heat dissipated by thermal conduction or to compensate

for variations in heating efficiency and perfusion, by estimating tissue parameters using MR

temperature measurements in a pre-treatment heating test [289, 292, 293] or during treatment

[294, 295]. Controllers designed for thermal ablation have also used adaptive approaches with

thermal dose as a control parameter to minimize treatment time [296] or avoid normal tissue

heating [297]. However, a comparison of fixed and adaptive control algorithms showed no

significant advantage for adaptive control in simulations of ultrasound heating with linearly

increasing perfusion, and only modest improvements in rise time and steady-state error in ex vivo

experiments [295].

Another variation of PID control defines an ideal energy distribution pattern to be applied

and divides it into a physically realizable series of focal point positions and energy deposition

parameters [298-300]. In early work a slow-moving hydraulic positioning system was used to

mechanically steer an ultrasound transducer along a spiral trajectory, building up to a desired

temperature distribution across a 10-20 mm diameter circle by the end of each 2-5 minute spiral

traversal [298, 299]. Recently, electronic focal steering has provided the ability to prescribe and

deposit energy at a series of points covering the entire three-dimensional target volume during

each MR thermometry image acquisition period [300]. In one report, the focus of a phased-array

transducer was electronically repositioned through a series of 70 ms sonications at up to 34

discrete locations in the 2.4 seconds between MRI temperature measurements, with the spatial

heating pattern customized every time a new set of images was acquired. However, nearly all of

the optimally-chosen sonication points ended up being located at the boundary of the targeted

region, the interior heated sufficiently by conduction [300]. This observation is consistent with

uniformity achieved by the peripheral energy distributions created with earlier mechanically

scanned focused ultrasound systems [216, 283].

Review of these reports suggests that mild hyperthermia can be delivered using a rapidly

scanned, mechanically-steered focused ultrasound transducer under MRI temperature control.

Using appropriately tuned fixed-gain proportional-integral controllers at several points along a

peripheral scan path can provide the spatially and temporally uniform temperature elevations

41

required to achieve localized drug delivery using thermosensitive liposomes. Mechanical

steering of a fixed ultrasound focus is a cost-efficient approach for approximating electronically-

steered clinical focused ultrasound systems, providing flexibility in developing customized

scanning trajectories and temperature control algorithms in preclinical investigations of image-

guided drug delivery.

1.5 Thermally mediated drug delivery using MRI-controlled focused ultrasound

Cancer is a heterogeneous evolutionary disease that commandeers normal cell mechanisms to

ensure its own survival. The clinical efficacy of chemotherapeutic agents is limited by systemic

toxicity and the inability to deliver a cytotoxic concentration to all cancer cells. Traditional

anticancer drugs are toxic to all rapidly dividing cells and have limited specificity for cancer

cells, whose rapid mutation rate allows drug-resistant subpopulations to emerge. Even if

molecular and cellular resistance is overcome, solid tumours present several barriers to the

delivery of potentially lethal drug concentrations to cancer cells.

Liposomes and other drug delivery systems improve drug targeting to tumours by taking

advantage of their increased blood vessel permeability and decreased nanoparticle clearance

from the tumour interstitial space. However, the clinical efficacy of liposomes has been limited

due to the sensitivity of passive liposome accumulation to the heterogeneity of vessel

permeability in tumours, poor nanoparticle transport within the tumour interstitium, and slow

drug release after accumulating in the tumour, in many cases preventing the delivery of lethal

concentrations of bioavailable drug.

Thermosensitive liposomes provide a mechanism for triggering the rapid release of high

concentrations of chemotherapeutic agents in a targeted region. Thermally mediated drug release

takes advantage of the chemosensitizing effects of hyperthermia, as well as increasing blood

flow to the tumour and increasing the extent and uniformity of tumour vessel permeability.

Rapid drug release from commercial thermosensitive liposome formulations allows release to

occur within the tumour vasculature, thus serving as a continuous intravascular infusion of free

42

drug originating at the tumour site. However, this localized drug release requires precise heating

to leverage the effects of hyperthermia on drug delivery and efficacy in tumours while

minimizing drug exposure in normal tissue. Preclinical studies have shown large increases in

tumour drug concentration by a proposed intravascular drug release mechanism, and there is

promising clinical evidence, despite limitations of clinical hyperthermia techniques.

Focused ultrasound is capable of noninvasively heating millimeter-sized regions deep

within the body, and can be combined with MR thermometry for precise temperature control

capable of achieving the tissue-specific heating necessary to localize drug delivery while

avoiding systemic drug release and normal tissue toxicity (Figure 1.11). Combining localized

hyperthermia under automatic feedback control with thermosensitive liposomes containing

Figure 1.11: Localized drug delivery using thermosensitive liposomes and MRI-controlledfocused ultrasound. MR temperature maps acquired rapidly during energy depositionusing externally focused ultrasound transducer (right). Temperature measured at focusused to adjust ultrasound energy deposition parameters in a feedback control loop(arrows), maintaining localized mild hyperthermia in a targeted region. Localizedhyperthermia provides continuous intravascular release of active drug in the targetedregion to drive drug diffusion into the tumour (left).

43

anticancer drugs could provide a noninvasive means of enhancing the effects of chemotherapy in

either a neoadjuvant role prior to surgery or in a curative approach for patients with unresectable

solid tumours.

1.6 Specific Aims This thesis describes the development of specific strategies for using MRI-controlled focused

ultrasound to achieve localized hyperthermia with precise spatial and temporal temperature

control, for the purpose of achieving image-guided thermally triggered drug release using

thermosensitive drug carriers. The ability of this approach to achieve targeted drug delivery and

improve drug distributions in the tumour microenvironment is evaluated.

The primary goals of this thesis are:

1) To develop clinically applicable strategies for using focused ultrasound and MR thermometry

to achieve precisely controlled, spatially and temporally uniform mild hyperthermia in a variety

of tissue types.

2) To study the effect of localized hyperthermia on thermally mediated drug delivery using

temperature sensitive drug carriers, as well as the mechanisms by which it occurs.

These goals were addressed in the following series of studies:

1. Development of a preclinical system to use MRI-controlled focused ultrasound hyperthermia

for localized drug delivery

The goals of this first study were to use MRI-controlled focused ultrasound to achieve precise,

noninvasive heating in vivo, and to demonstrate the feasibility of localized drug release in

normal tissue using temperature sensitive drug carriers and MRI-controlled focused ultrasound

hyperthermia. In this study, a preclinical MRI-controlled focused ultrasound system developed

as a platform for performing studies of controlled hyperthermia and thermally mediated drug

delivery in rabbits was described, and the feasibility of using ultrasound hyperthermia to achieve

44

localized drug release from thermosensitive liposomes was demonstrated in normal rabbit

muscle.

2. Study of localized drug delivery: applications in bone

After demonstrating the feasibility of achieving localized drug release in homogeneous tissue, a

second study evaluated the potential of combining MRI-controlled ultrasound hyperthermia with

thermosensitive liposomes for the treatment of bone metastases. Targeting the femur in healthy

rabbits, strategies for using MR thermometry to control FUS heating at the bone interface based

on MR temperature measurements in adjacent soft tissue were described, and the use of

ultrasound to heat through the bone to achieve localized drug delivery was demonstrated.

Numerical simulations estimated cortical bone temperatures for two different heating strategies.

Significant increases in doxorubicin concentration in heated vs. unheated muscle and marrow

were observed. These results suggest that bone heating can be safely controlled using

temperatures measured in adjacent soft tissue, and that this can be used to achieve locally

enhanced doxorubicin deposition in soft tissues within and surrounded by bone.

3. Drug delivery in a rabbit model of solid tumours: biodistribution and drug distribution in the

tumour microenvironment

The first two studies demonstrated that the combination of heat and thermosensitive liposomes

enhances overall drug accumulation by 10-20 times in homogenized tissue samples. However,

those measurements of the overall quantity of doxorubicin in heated regions provide no

information about the distribution of doxorubicin in heated regions, or the ability to reach all

targeted cells with a cytotoxic dose. Also, these studies were done in normal tissue with largely

homogeneous heating and perfusion, and do not address the specific challenges of drug delivery

in the tumour microenvironment.

In the third study, fluorescence microscopy was employed to verify that this increased

overall accumulation corresponds to high levels of bioavailable drug getting to their active sites

in the nuclei of tumour cells. The large size and heterogeneous vasculature of the rabbit VX2

tumour model permits testing of the robustness of heating techniques, and provides insight on

45

drug delivery in heterogeneous solid tumours. This study extends previous work by

demonstrating uniform temperature control in large rabbit VX2 tumours, and by investigating the

microregional distribution of released bioavailable drug within the tumour microenvironment.

The results of this thesis demonstrate image-guided drug delivery in a large animal model

using a commercially-developed liposome formulation and clinically applicable heating

strategies. MRI-controlled focused ultrasound is a noninvasive means of achieving precise,

localized heating in a variety of deep-seated tissues, and can be used with temperature-sensitive

drug carriers to achieve large increases in drug deposition in soft tissue, at bone interfaces, and in

solid tumours. We show that following heat-triggered release, bioavailable drug accumulates in

the nuclei of tumour cells where it can exert cytotoxic effects. These results will be useful in the

development of clinical strategies for image-guided drug delivery using ultrasound hyperthermia

and temperature-sensitive drug carriers for a variety of tumour types.

47

Chapter 2 Localized drug release using MRI-controlled focused ultrasound hyperthermia1

2.1 Introduction Hyperthermia has been shown to enhance extravasation of drug carriers in solid tumours,

potentially overcoming barriers to drug delivery presented by the tumour microenvironment [85].

The use of temperature-sensitive drug carriers provides a mechanism for triggering the rapid

release of high concentrations of active drug in a targeted region [85, 168, 301-303]. This

localized therapy depends on the stability and heat-triggered release of the contents of drug

carriers, and also the ability to accurately maintain controlled temperature elevations in targeted

regions with minimal temperature increase in surrounding tissue.

Most studies with thermosensitive drug carriers in preclinical models have either applied

regional, superficial heating with water baths [76, 84, 129, 130, 173, 183, 301] and microwave

electromagnetic applicators [62, 181], or localized heating of deep targets with invasive catheters

[77, 79, 147]. Recently, pulsed high-intensity focused ultrasound has been used to potentiate

drug release from thermosensitive liposomes [68, 178, 304]. Temperature measurement in all of

these studies has been limited to point measurements made by invasive thermocouple or optical

probes.

Millimetre-sized focal regions several centimetres beneath the skin can be heated

noninvasively with minimal energy deposition in the intervening tissue using externally focused

ultrasound (FUS) transducers [305]. By mechanically [211, 220] or electronically [306] scanning

the ultrasound focus, larger target regions can be heated. Early mechanically scanned systems 1 This chapter is adapted from the article: “Localised drug release using MRI-controlled focused ultrasound hyperthermia”. Robert Staruch, Rajiv Chopra, Kullervo Hynynen. International Journal of Hyperthermia 27(2): 156-171 (2011).

48

[220] overcame spatially varying energy absorption and blood flow in tissue with rapid scanning

of the focal point [212, 219] and multi-point control based on arrays of invasive thermocouple

probes [258, 259]. Promising clinical temperature distributions were achieved in a number of

anatomical sites [307].

Since then it has been shown that spatial temperature distributions can be noninvasively

monitored in most tissues using MRI thermometry based on the proton resonance frequency shift

[261, 267]. Stable measurements can be made during ultrasound heating [226], and have been

used clinically to guide point-by-point coagulation of tumours [229, 234, 238, 240]. MRI

temperature measurements can also be used to regulate temperature elevations in a feedback

control loop [289, 292, 293, 296]. This form of adaptive, closed-loop therapy has been applied

for tumour coagulation using intracavitary [290, 308] and externally focused [309] ultrasound

transducers, lasers [310], and percutaneous radiofrequency applicators [311]. Noninvasive

radiofrequency hyperthermia under MRI temperature control is under development [312], but

poor steering and focusing abilities limit target temperature elevations that can be safely

achieved. MRI-controlled hyperthermia using scanned ultrasound beams is challenging to

implement due to image artifacts caused by device motion and the inability to use conventional

motors in the magnet bore. Slow hydraulic scanning of a single element transducer has been

shown to achieve spatially uniform heating across a 20 mm diameter region under MRI control,

but with wide periodic temporal variations in temperature elevation [298, 299]. With a

sophisticated phased-array system capable of rapid electronic displacement of the focal point,

precise spatial and temporal control has been achieved [300]. An MRI-compatible focused

ultrasound system developed recently [313] enables mechanical scanning of an ultrasound

transducer at circular scan rates required to achieve uniform temperature elevations in tissue (>

0.2 Hz) [212], taking advantage of simple multi-point control schemes previously used under

thermocouple control [258, 259].

The objective of this study was to investigate the use of MRI-controlled focused

ultrasound hyperthermia to achieve localized drug release in vivo. For this preliminary feasibility

study, a rabbit model was used. Normal thigh muscle was heated using a mechanically-scanned

focused ultrasound system under MRI temperature control, and temperature-sensitive liposomes

49

were administered during heating. The ability to maintain spatially and temporally uniform

hyperthermia in 10-15 mm diameter regions was evaluated, as well as the ability to achieve

localized accumulation of released drug.

2.2 Materials and Methods

2.2.1 MRI-controlled focused ultrasound system

Noninvasive heating of rabbit thigh was achieved by continuous sonication using a spherically

focused, air-backed piezoceramic ultrasound transducer (fundamental frequency, 0.536 MHz;

curvature radius, 10 cm; aperture diameter, 5 cm) driven at its fifth harmonic (2.787 MHz). To

achieve spatially uniform temperature distributions in the thigh under MRI temperature feedback

control, the transducer was mounted to a custom-made MRI-compatible three-axis positioning

system [313] programmed to move along a 1-2 cm diameter circular trajectory at a speed of 1

revolution per second, for simultaneous ultrasound heating and MRI thermometry in the plane of

the ultrasound focus. The positioning system was constructed from non-magnetic components,

piezoceramic actuators and optical encoders, and was capable of simultaneous transducer motion

and MR imaging with minimal mutual interference. The system was able to achieve a spatial

positioning precision of approximately 0.1 mm with linear speeds up to 50 mm/s per axis over a

total travel of 4 cm in each direction.

The transducer diameter (5 cm) and focal length (10 cm) were chosen to reduce magnetic

susceptibility artifacts caused by the motion of the transducer during heating. Driving the

transducer at 2.787 MHz (fifth harmonic) resulted in a focal spot that was 1.4 mm wide by 20.7

mm long, centered 94 mm from the transducer surface (full-width half-maximum of the relative

pressure amplitude squared, measured in water with a fiber-optic hydrophone). This spans the

length of preclinical tumours of interest with minimal energy deposition beyond the focal point,

but allows for modulation of heating in the transverse plane. The transducer was driven by an

arbitrary waveform generator (33250A, Agilent, USA) and radiofrequency power amplifier

(NP2912, NP Technologies, Inc., Newbury Park, CA). The electrical power was measured with a

power meter (438A, Hewlett Packard, Palo Alto, CA) connected to the forward and reverse

50

signal of a dual directional coupler (C1373, Werlatone, Brewster, NY). The transducer’s

electrical impedance was matched to the output impedance of the amplifier (50 Ω) with a

custom-made passive matching circuit. At the driving frequency, the emitted acoustic power was

approximately 70% of the applied electrical power (forward minus reflected) when measured

with a radiation force balance technique [314] using a laboratory balance (AE200, Mettler-

Toledo Inc., Columbus, OH). This efficiency was used to calculate the electric power required to

achieve a desired acoustic power during sonication.

2.2.2 Animals

The experiments described in this study were approved by the institutional Animal Care

Committee at Sunnybrook Health Sciences Centre. New Zealand White rabbits (3-4 kg, n=10)

were anaesthetized with an intramuscular injection of 50 mg/kg ketamine and 10 mg/kg xylazine

administered away from the planned sonication site, before intubation and maintenance under 2-

4% isoflurane. To prevent motion, an additional 1.0 ml injection of 3 parts ketamine/xylazine

diluted with 7 parts saline was administered into a cannulated ear vein approximately 5 minutes

prior to hyperthermia.

For this study, a rabbit model was selected in order to provide a large enough tissue mass

for targeting, and to reduce the effect of prolonged localized heating on core body temperature

that might be seen in smaller animal models. Anaesthetized rabbits had their thighs shaved and

depilated, and were placed on a stage above the focused ultrasound system in the lateral

decubitus position, as illustrated in Figure 2.1A. Degassed water was used to couple the target

thigh through an 80 mm square window of 25 μm polyimide film into the reservoir of degassed

water in which the transducer was submerged. A square (87 mm inner side length) custom-made

single-loop RF receive coil with a square opening in the middle (85 mm inner side length) was

designed to fit underneath the animal around the window for optimal SNR in the heated region.

The thigh was positioned such that the ultrasound beam made approximately normal incidence

with the skin through the coil. The transducer position was adjusted to locate the focus at a depth

of 1-2 cm from the skin, at the interface between the biceps femoris and the semimembranosus

or semitendinosus muscles of the thigh. Once the correct focal depth was achieved, the

51

Figure 2.1: Experimental setup for in vivo MRI-controlled focused ultrasoundhyperthermia. A) Axial depiction of rabbit in lateral decubitus position. B) Axial T2-weighted image along beam path, indicating the coronal plane used for MRI thermometryand the acoustic field path at the boundaries of the heating trajectory. C) Coronal T2-weighted image taken through the plane indicated in B, with eight control regions shownon the periphery of the 10 mm diameter target region. B, C have 120 mm field of view.

52

transducer was scanned along a circular trajectory transverse to the ultrasound beam direction.

The diameter of the circular trajectory was varied between 1.0 and 1.5 cm across experiments

based on the available acoustic window. To prevent undesired tissue heating due to acoustic

reflections and propagation into the contralateral thigh, a saline bag and polyurethane rubber

acoustic absorber (AptFlex F28, Precision Acoustics, Dorchester, Dorset, UK) were placed

between the rabbit’s thighs, coupled with ultrasound gel. Thermosensitive liposomal doxorubicin

and MR contrast agents were injected into a cannulated ear vein.

The rabbit’s body temperature was monitored by a fibre-optic temperature probe inserted

into the rectum (3100, Luxtron Corp., Santa Clara, CA), and maintained by regulating the

temperature of a hot water blanket covering the animal (Gaymar T/Pump Model TP-500,

Gaymar Industries, Inc., Orchard Park, NY). The temperature of the degassed water reservoir

below was maintained by pumping warm water through a heat exchanger made from a coil of

polyurethane tubing. These temperatures were monitored with additional fibre-optic probes at the

skin/blanket interface and in the water bath. Oxygen saturation (SpO2) and heart rate were

monitored by a veterinary pulse oximeter (8600V, Nonin Medical, Inc., Plymouth, MN).

In early experiments, discrete movements of 2-3 mm were likely the result of bone

heating caused by transmission of ultrasound through the animal to the contralateral thigh, as

well as acoustic reflection at skin/air interfaces. In subsequent experiments motion was avoided

by careful positioning to avoid bone, and by an intramuscular injection of local analgesic (2%

Lidocaine HCl Injection USP, Alveda Pharmaceuticals, Toronto, ON) immediately before

hyperthermia, administered upstream from the planned sonication site.

In one of the treated animals, oxygen saturation and heart rate dropped suddenly from

99% to 50% and from 150 to 100 beats per minute shortly after drug injection during heating,

and required resuscitation before post-treatment imaging. These symptoms are consistent with

liposome-related anaphylactoid reactions reported previously in dogs [181], pigs [315] and some

humans [316], and could likely be prevented by premedication with steroids and antihistamines

[181, 315].

53

2.2.3 MR imaging

Experiments were performed with the positioning system placed in the bore of a clinical 1.5T

MR imager (Signa, GE Medical Systems, Milwaukee, WI). Co-registration of the ultrasound

focus in the MRI coordinate system was accomplished by first heating a gel phantom and

locating the region of temperature increase with MRI thermometry.

Next, the animal was positioned on the FUS system in the magnet, and axial and coronal

anatomical T1-weighted (FSE-XL, TE = 15 ms, TR = 500 ms, ETL = 2, BW = 15.63 kHz, NEX

= 3, 256 x 128 matrix, FOV = 12 cm, slice = 3 mm, no phase wrap) and T2-weighted (FSE-XL,

TE = 75 ms, TR = 2000 ms, ETL = 4, BW = 15.63 kHz, NEX = 2, 256 x 128 matrix, FOV = 12

cm, slice = 3 mm, no phase wrap) images were acquired to locate the target region in the thigh

before heating (Figure 2.1B, C). These images were re-acquired after heating to evaluate changes

in tissue perfusion and/or tissue damage caused by ultrasound heating. The post-sonication T1-

weighted images were acquired before and continuously for 5 minutes after administration of a

2-second bolus ear vein injection of MR contrast agent (0.2 mmol/kg, Omniscan, GE Healthcare)

followed by a 1 ml injection of saline.

During mechanical scanning of the ultrasound transducer, fast spoiled gradient-echo

images (FSPGR, TE = 10 ms, TR = 38.6 ms, 30° flip angle, BW = 31.25 kHz, 128 x 128 matrix,

FOV = 12 cm, slice = 5 mm, NEX = 1, acquisition time = 5 s) were continuously acquired in a

coronal plane through the acoustic focus to measure the spatial heating pattern in the target

region. A real-time image acquisition interface [317, 318] was used to transfer raw k-space data

to a control computer immediately after acquisition for image reconstruction and processing of

temperature maps using software written in MATLAB (Mathworks, Natick, MA). Temperature

maps were calculated by performing a complex phase subtraction between treatment images and

the average of five initial baseline images acquired prior to heating, scaling the result to a

relative temperature using a proton resonance frequency shift coefficient of -0.010 ppm/°C [268],

and then adding the baseline temperature measured by a fibre-optic temperature probe in the

rectum. Unheated regions of the FSPGR images were identified to measure phase drift over the

experiment duration.

54

Accurate measurement of temperature from the phase difference between MR images

acquired during mechanical scanning of an ultrasound transducer is a significant technical

challenge. A moving object in the bore of the MRI causes a local change in the main magnetic

field related to its magnetic susceptibility [319] violating the assumption of static field

inhomogeneities in the PRF shift subtraction method [261]. The circular scan trajectory of the

ultrasound transducer was found to cause periodic phase shifts, whose frequency was related to

the speed of rotation and whose amplitude increased with proximity of the transducer to the

imaging plane. As temperatures were expected to be relatively stable during treatment, the

resulting temperature artifacts were mitigated by temporal averaging. Tests of several window

lengths indicated that by averaging over a 30 second sliding window, the standard deviation of

temperatures measured during transducer motion could be matched to that seen without motion.

This window was applied during treatment to temperature images used for feedback control, as

well as for images used to calculate thermal dose.

2.2.4 Closed-loop feedback control

Previous studies [259, 300] suggest that single-point fixed-gain proportional-integral (PI) control

applied at multiple points along a rapidly scanned peripheral energy deposition trajectory can be

used to achieve uniform temperatures in a targeted region with sharp gradients at the boundaries

and a reasonable degree of steady-state temperature accuracy, robust to spatial and temporal

changes in blood flow.

In this study, spatial heating was controlled by eight independent PI feedback controllers

along the scanned heating trajectory, rapidly switching the applied power as the transducer

scanned along the circular trajectory, and updating the controller outputs after each image

acquisition according to the following equation:

[ ] [ ]( ) [ ]( )∑ −+−= igoalIngoalPn tTTKtTTKtP 8..18..18..1 [Eq. 6]

For temperature image n, the temperature T in each of the eight pre-defined 3 x 3 pixel control

regions was compared with the desired target temperature Tgoal to calculate an error term. This

error term was scaled by proportional gain KP, while the integral of the error over all previous

55

images was scaled by integral gain KI to specify a new acoustic power P to be delivered by the

transducer as the focus crossed that region. When the transducer position coincided with one of

the control locations during scanning, the function generator output amplitude was adjusted to

achieve the acoustic power determined by the control algorithm. This modulation of function

generator amplitude occurred at a frequency of 8 Hz during circular scanning at 1 Hz and

temperature imaging at 0.2 Hz. Applied power was limited to 16 acoustic watts, corresponding to

a measured peak negative pressure of 2.5 MPa at the focus, safely below the predicted threshold

for cavitation to occur at this frequency (0.6 MPa + 5.3 MPa MHz-1, or 15.37 MPa at 2.787

MHz) [320].

During gel and in vivo heating experiments, eight equally-spaced independent control

points were selected around the target trajectory with the controller gains identified in

simulations, as described below. Circular transducer motion and then continuous FSPGR image

acquisition were initiated. After five reference FSPGR images were acquired and averaged,

twelve additional images were acquired to determine the standard deviation of temperature

measurements before ultrasound heating was initiated under temperature feedback control with

the positioner rapidly scanning the focus along a circular trajectory. In early experiments,

periodic temperature artifacts and proportional-integral control of voltage instead of power (thus

varying with the temperature error squared) caused excessive temperature fluctuations in the

heated region. Averaging control point temperatures over a 30 second sliding window and using

a less aggressive power-based controller (KP = 4.5, KI = 0.03) that tracked an exponential rather

than step input function gave accurate temperature control with steady-state oscillations of less

than 1°C and no overshoot.

To evaluate the temporal and spatial uniformity of heating, the mean, T90 (temperature

that 90% of the target exceeds) and T10 (temperature that 10% of the target exceeds)

temperatures were measured at each time point across circular regions of interest matching the

intended target diameter. The steady-state mean and standard deviation of each parameter was

calculated over all images starting once Tgoal reached 95% of the final target temperature

elevation. Thermal dose in the target region was calculated in cumulative equivalent minutes at

43°C (CEM43) using the Sapareto and Dewey time-temperature equation [103].

56

2.2.5 Numerical Simulations

Numerical simulations of scanned ultrasound heating were used to identify appropriate controller

gain values as in previous simulation studies of scanned focused ultrasound heating [212, 259].

Briefly, a multilayer transmission model of the acoustic field pattern [321, 322] with the focal

plane at 3 cm into a 5 x 5 x 7 cm block of homogeneous, non-perfused tissue was shifted off-

center by the scan radius and moved in a circular path around the center using affine

transformations and linear interpolation. At each 0.01 second time step, the three-dimensional

temperature distribution was recalculated by the Pennes bioheat transfer equation [214] using a

previously described finite-difference time-domain implementation [322]. The simulations

assumed a tissue density of 1000 kg m-3, specific heat capacity of 3700 J kg-1 °C-1, thermal

conductivity of 0.5 W m-1 °C-1, amplitude attenuation coefficient of 4.1 Np m-1 MHz-1, and speed

of sound of 1570 m s-1, with a grid size of 0.25 x 0.25 x 1.0 mm and boundary conditions of

water at 37°C on all surfaces. To model feedback control, temperatures on the simulation grid

were averaged spatially and temporally, and zero-mean Gaussian noise with a standard deviation

of 0.5 was added to match the resolution, sampling rate and background temperature

measurement noise of MR thermometry (1 x 1 x 5 mm, 5 seconds, ±0.5°C). Controller settings

that appeared promising in simulation were tested in a tissue-mimicking gel made primarily of

agar and condensed milk [323], under similar conditions to those used for in vivo experiments.

2.2.6 Sonication, drug administration and tissue harvesting

MRI-controlled focused ultrasound was used to heat 10 to 15 mm diameter regions within the

thigh to 43°C for a minimum of 20 minutes. The temperature elevation and exposure duration

were chosen to be sufficient to trigger drug release from temperature-sensitive liposomes [301]

without causing significant tissue damage [106].

Pegylated lyso-thermosensitive liposomal doxorubicin (LTLD, Thermodox®, Celsion

Corporation, Columbia, MD) was provided by the manufacturer at a concentration of 2 mg

doxorubicin/ml. Rabbits were prescribed a doxorubicin dose of 2.5 mg/kg, administered as a

dilution of LTLD in 5% dextrose sterile solution to an injection volume of 10 ml, infused slowly

57

into the ear vein during MRI-controlled focused ultrasound heating. Once the average

temperature in the target region reached 43°C (approximately 6 minutes), the 10 ml infusion

volume was administered over 8 minutes at 1.2 ml/min using an MRI-compatible injection

system (Spectris Solaris EP, MEDRAD, Inc., Warrendale, PA). The dose used in these

experiments corresponds to a human equivalent dose of approximately 30 mg/m2 [324], which is

similar to that used in dogs (14-20 mg/m2 [181]) and rats (30 mg/m2 [79]), and on the low end of

what has been administered in clinical trials.

Upon completion of heating, image acquisition continued until target temperatures

returned to resting temperature to quantify the total thermal dose accumulated in the target region

and surrounding tissues. Post-treatment T2-weighted and contrast-enhanced T1-weighted images

(0.2 mmol/kg, Omniscan, GE Healthcare) were acquired to assess changes in tissue perfusion

and/or tissue damage caused by ultrasound heating. After imaging, skin in the beam path was

examined for burns. Trypan Blue (2% weight/volume) was administered intravenously at 1.5

ml/kg; increased extravasation of the dye was used to identify sonicated regions during

dissection, based on increased vascular permeability in heated tissue.

Approximately two hours after LTLD infusion, unabsorbed liposomal drug was flushed

from the vasculature by transcardiac perfusion with saline. After sacrifice, 25-50 mg tissue

samples placed in 1.5 ml microcentrifuge tubes were snap-frozen in liquid nitrogen and stored at

-80°C. Several samples of muscle tissue were collected from both the heated region of the treated

thigh, as well as unheated regions in both the treated and untreated thighs to evaluate the spatial

heterogeneity of drug release.

2.2.7 Analysis of drug concentration and release

Tissue doxorubicin concentrations were measured by the fluorescence intensity of doxorubicin in

homogenized tissue samples using published extraction techniques [58, 62, 68, 70, 325]. Upon

removal from storage, tissue samples were pulverized with pestle and mortar immersed in liquid

nitrogen. The resulting powder was weighed and added to 20 volumes of acidified ethanol

extraction solvent (0.3N HCl in 50% ethanol) before homogenization with a pestle and tube

58

grinder (KONTES® Tenbroeck, Kimble Chase LLC, USA), and overnight refrigeration in

acidified ethanol solvent. The next day, homogenates were centrifuged at 16000 x g for 30

minutes at 4°C using a benchtop centrifuge (Centrifuge 5417C, Eppendorf, Hamburg, Germany).

Care was taken to pipette as much of the clear supernatant as possible into a fluorometry cuvette,

but the amount varied, primarily by the mass of the sample. Acidified ethanol was added to make

a standard volume of 2 ml, matching the volume used for calibration. A benchtop fluorometer

(VersaFluor, Bio-Rad Laboratories, Hercules, CA) was used to measure the fluorescence

intensity of doxorubicin in the extracted supernatant, recorded as the average of three readings

taken with a 480 nm excitation and 590 nm emission filter set. Relative fluorescence intensities

were scaled to doxorubicin concentrations by comparison with a calibration curve made from the

fluorescence intensities of a serial dilution of free doxorubicin (Doxorubicin HCl, Teva

Novopharm, Toronto, ON) in acidified ethanol, normalized by the mass of the sample, and

corrected for variations in added extraction fluid volume. When known amounts of doxorubicin

were added to thigh muscle from untreated rabbits before homogenization, this procedure

resulted in extraction efficiencies of approximately 80%. Drug concentrations are reported in this

study without adjustment.

The Wilcoxon matched pairs signed rank test was used to compare the difference in drug

concentrations between tissues samples from heated and unheated regions of normal thigh

muscle without assuming that the data are normally distributed. Differences were considered

statistically significant for values of P less than 0.05.

2.3 Results

2.3.1 Simulations and in vitro temperature control

Figure 2.2A shows temperature elevations in a 1 cm diameter circular target for simulations of

proportional-integral control using proportional gain KP = 4.5, integral gain KI = 0.03, and an

exponential input function with a target temperature elevation of 10°C and time constant τ = 160

seconds. This time constant was chosen to give a rise time long enough that the input could be

tracked at each control point, thus preventing an initial accumulation of error and allowing the

59

use of integral smoothing without overshoot. Temperature control using the same controller

gains and exponential time constant to create a 10°C temperature elevation in a tissue-mimicking

gel phantom is shown in Figure 2.2B, demonstrating good temporal and spatial control (time to

reach 95% of target temperature elevation = 365 seconds, steady-state mean ± temporal standard

deviation = 10.0 ± 0.3°C, T90 = 9.3°C, T10 = 10.6°C). Similar performance was observed in

simulation (rise time = 515 seconds, mean = 9.9 ± 0.2°C, T90 = 9.7°C, T10 = 10.1°C).

Temperature maps taken at 10 minutes after the start of heating are shown in Figure 2.2C-D,

demonstrating spatially uniform heating. Thermal dose maps calculated from these simulation

and gel experiments, assuming a target temperature of 43°C, are shown in Figure 2.2E-F. These

results indicate that uniform heating was maintained over time to produce a uniform cumulative

effect, using multi-point control robust to either normally-distributed simulated noise or periodic

MRI temperature measurement fluctuations caused by transducer motion during heating.

2.3.2 MRI-guided focused ultrasound hyperthermia in vivo

The spatial location of the heated region and the eight ROIs used for feedback control are shown

for rabbit #9 in the coronal FSPGR magnitude image in Figure 2.3A, acquired in the same plane

as the T2-weighted image in Figure 2.1C. The target was prescribed as a 10 mm diameter circle,

centered 2 cm from the skin. Temporal evolution of the spatial temperature distribution in this

plane during heating is shown in Figure 2.3B-D with temperatures greater than 37°C overlaid

upon the corresponding magnitude images at selected time points of five, ten, and twenty

minutes after heating began. Following an initial period of peripheral temperature elevation, a

relatively uniform region of spatial heating was observed, with a rapid falloff outside the target

region.

The temperature elevation in a circular ROI covering the 10 mm diameter target region is

shown in Figure 2.4A for the same experiment. At each time point, the mean ROI temperature,

the temperature which 90% of the ROI exceeds (T90), and the temperature which 10% of the ROI

exceeds (T10) are shown. Also shown are the mean temperatures in unheated regions 1.0 to 1.25

cm from the edge of the circular target, the controller input function Tgoal, and the start and end

times of the LTLD infusion. The observed temperature increase of approximately 1°C in the

60

Figure 2.2: Simulations (A,C,E) and in vitro (B,D,F) testing of proportional-integralMRI temperature control for mechanically-scanned focused ultrasound hyperthermia.Mean, T90 and T10 temperatures in the 10 mm diameter circular region of interest for atarget temperature elevation (Tgoal) of 10°C using controller gains KP = 4.5, KI = 0.03 andan exponential input function with time constant τ = 160 seconds, in simulation (A) and invitro, heating a tissue-mimicking gel phantom (B). Temperature maps at 10 minutes afterthe start of heating, and thermal dose maps calculated from the same simulation (C,E) andin vitro (D,F) controlled heating data, assuming a target temperature of 43°C. 10 mmdiameter circular scan trajectory outlined in white.

61

unheated region represents a combination of heat conduction from the target region, variations in

water bath temperature during treatment, and possible phase drift. Transducer motion caused

periodic artifacts in raw FSPGR phase images that resulted in a temperature measurement

standard deviation of approximately ±1°C in pre-treatment images across the target ROI, reduced

to ±0.4°C when a 6-image windowed average was applied during treatment. After averaging, the

mean temperature in the ROI over the treatment interval was 43.2 ± 0.3°C after a rise time of 415

Figure 2.3: Evolution of temperature distribution during MRI-controlled focusedultrasound heating of a 10 mm diameter region. Control ROIs, circular unheated ROI, andtemperatures greater than 37°C are overlaid on coronal FSPGR magnitude images after(A) 0 (B) 5, (C) 10, and (D) 20 minutes of heating.

62

seconds, with an average T90 and T10 of 41.7°C and 44.6°C, respectively. The thermal dose is

shown as the number of cumulative equivalent minutes at 43°C (CEM43°C) in Figure 2.4C, and

in this case had a median of 25 CEM43°C in the ROI, with CEM43°C,T90 and CEM43°C,T10 of

4 and 65 CEM43°C, respectively. Insight on the large variability in thermal dose can be gained

from the mean ± standard deviation of temperature versus radius shown in Figure 2.4B, which

suggests that variations result from consistently higher temperatures in the center of the heated

Figure 2.4: Example of temporal and spatial temperature control over a 10 mm diametercircular area using MRI-controlled focused ultrasound hyperthermia. A) Mean, T90 andT10 temperatures in the circular region of interest. Lyso-thermosensitive liposomaldoxorubicin (LTLD) infusion duration is shaded. At the end of heating, the transducerreturned to the center position and reduced temperature noise was observed. B) Time-averaged mean ± SD temperature versus radius. Unheated region of interest is shaded. C)Thermal dose map centered on target region, circular scan trajectory outlined in white.

63

region rather than temporal fluctuations at a given point. The variation between CEM43°C,T90

and CEM43°C,T10 is further increased by the 43°C breakpoint defined in the thermal dose

calculation: four minutes at 42°C contributes the same thermal dose as one minute at 43°C.

Spatial non-uniformity could be decreased by modifying the controller to force output power to

zero if any measured temperature exceeds a certain value [259], or by increasing the scan radius

to reduce thermal buildup in the center of the scan trajectory [212, 286]. This latter method was

used in later experiments to avoid excessive thermal exposures in the central region.

Table 2.1: Summary of in vivo MRI-controlled FUS experiments for hyperthermia-mediated drug delivery.

Rabbit Target Radius

Mean ± SD temperature

Mean ± SD T90

Mean ± SD T10

Median thermal

dose

Mean ± SE Doxorubicin in

unheated muscle

Mean ± SE Doxorubicin in heated muscle

(mm) (°C) (°C) (°C) (CEM43) [ng/mg] [ng/mg]

Controlled hyperthermia (n = 6)

1 5.0 43.7 ± 1.1 41.2 ± 1.0 46.1 ± 1.4 68.9 0.2 ± 0.0 4.0 ± 0.8 2 5.0 42.1 ± 0.4 39.2 ± 0.6 45.2 ± 0.9 8.8 0.5 ± 0.1 9.6 ± 3.2 5 5.0 43.0 ± 0.4 41.7 ± 0.4 44.2 ± 0.4 26.3 0.4 ± 0.0 0.6 ± 0.0 7 5.0 43.9 ± 0.6 42.5 ± 0.6 45.2 ± 0.6 41.6 0.7 ± 0.2 14.3 ± 1.9 9 5.0 43.2 ± 0.3 41.7 ± 0.3 44.6 ± 0.3 25.4 0.3 ± 0.1 3.3 ± 0.4 10 7.5 41.7 ± 0.9 39.9 ± 1.0 43.4 ± 0.9 7.9 0.8 ± 0.2 18.1 ± 1.2

Mean ± SD 42.9 ± 0.6 41.0 ± 0.7 44.8 ± 0.8 29.8 0.5 ± 0.2 8.3 ± 6.9

Thermal coagulation (n = 4)

3* 7.5 - - - >240 0.3 ± 0.0 5.4 ± 3.6 4* 7.5 - - - >240 1.0 ± 0.1 11.9 ± 3.5 6* 5.0 - - - >240 0.6 ± 0.1 2.2 ± 0.5 8* 5.0 - - - >240 0.4 ± 0.0 2.0 ± 0.9

Mean ± SD - - - >240 0.6 ± 0.3 5.4 ± 4.6

* In experiments where MRI thermometry was corrupted by rabbit motion, sonication continued at high power and caused coagulation of the target region.

64

Hyperthermia experiments are summarized in Table 2.1, where the mean, T90 and T10

temperatures are reported as the average of each measure in the target ROI ± the standard

deviation over the time interval during which the target temperature was maintained. On average,

Figure 2.5: Tissue damage observed using T2-weighted and contrast-enhanced T1-weighted imaging following thermal coagulation in rabbit thigh. A, B) T2-weighted axial(A) and coronal (B) images through the center of the heated region. C, D) Correspondingcontrast-enhanced T1-weighted axial (C) and coronal (D) images from the sameexperiment. 80 mm field of view, targeted regions (outlined) and thermal damageboundaries (arrows) shown.

65

the steady state mean, T90 and T10 temperatures were 42.9°C, 41.0°C, and 44.8°C, respectively,

and the mean temperatures had a standard deviation over the treatment duration of ±0.6°C. In

four experiments, small amounts of motion (2-3 mm) caused temperature artifacts that varied

across the image between 2 and 10 degrees below the actual temperature. Continued closed-loop

sonication caused higher achieved temperatures that resulted in coagulative necrosis of the target

region, with an assumed thermal dose of greater than 240 CEM43°C. In these cases, tissue

damage was detectable as regions of low signal intensity surrounded by high-intensity rings on

post-treatment T2-weighted and contrast-enhanced T1-weighted images as shown in Figure 2.5,

and Trypan blue extravasation observed after cardiac perfusion was localized to the periphery of

coagulated regions. In all other experiments stable temperature imaging was achieved, and

rabbits received a median thermal dose of 28.9 ± 23.4 CEM43°C. In these cases, no changes

were observed on post-treatment imaging, and upon dissection little to no Trypan blue

extravasation was observed in the heated region.

2.3.3 Hyperthermia mediated drug delivery

The mean and standard deviation of doxorubicin concentration measured in tissue samples

collected from heated and unheated regions of normal thigh muscle from each of ten rabbits is

reported in Table 2.1. As summarized in Figure 2.6 across the six animals that received

controlled hyperthermia, drug concentrations in heated regions were on average 15.8 times

higher than in the corresponding unheated regions of the contralateral thigh (95% confidence

interval [CI] = 7.0 to 24.6 times). One animal in the hyperthermia group had minimal effect

(number 5). The pair-wise difference in doxorubicin concentration between heated and unheated

regions was statistically significant (8.3 versus 0.5 ng/mg, per-animal difference = 7.8 ng/mg,

95% CI = 0.8 to 14.9 ng/mg, P < 0.05, Wilcoxon matched pairs signed rank test). In the four

cases where thermal coagulation occurred, coagulated regions had 9.7 times higher doxorubicin

concentrations than the contralateral thigh (range = 3.7 to 18.8 times), but the pair-wise

difference did not reach statistical significance (5.4 versus 0.6 ng/mg, per-animal difference =

4.8 ng/mg, 95% CI = -2.2 to 11.8 ng/mg, P = 0.13).

66

In one rabbit, additional tissue samples were harvested near the heated region to

investigate drug deposition as a function of radius. Figure 2.7 shows tissue doxorubicin

concentrations at 0, 5, 10, and 20 mm from the centre of the 5 mm radius heated region, as well

as in the unheated contralateral thigh. Drug concentration in tissues sampled at 0, 5, and 10 mm

from the center of the treated region were higher than at 20 mm from the center of the treated

region, or in the contralateral thigh (3.3, 9.6, and 7.1 versus 0.5 and 0.3 ng/mg, respectively).

Figure 2.6: Doxorubicin concentrations measured by fluorescence intensity in tissuesamples harvested from heated and unheated regions of thigh muscle in rabbits receivingcontrolled hyperthermia (n = 6) or thermal coagulation (n = 4). Columns, mean; bars, SD.* P < 0.05, Wilcoxon matched pairs signed rank test.

67

2.4 Discussion

2.4.1 MRI-controlled focused ultrasound hyperthermia

In this study, temporally and spatially uniform localized hyperthermia was achieved in normal

rabbit thigh muscle using mechanically scanned focused ultrasound heating under multi-point

proportional-integral MRI temperature control. Mean target temperatures of 42.9°C were

maintained over approximately 20 minutes with a temporal standard deviation of 0.6°C and an

average T90-T10 of 3.8°C across the 10 to 15 mm diameter circular target region.

Figure 2.7: Doxorubicin concentrations measured by fluorescence intensity in tissuesamples from rabbit #9, harvested at 0, 5, 10, and 20 mm away from the center of theheated region, as well as from the unheated contralateral thigh. Overlay: time-averagedtemperature profile. Inset: sample positions shown on image of mean temperatures greaterthan 37°C.

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2.4.2 Sources of error in MRI thermometry

A major drawback to heating with a moving ultrasound transducer in an MRI is that as the

transducer moves along its circular scanning trajectory, the susceptibility distribution in the bore

changes, giving rise to periodic variations in phase images and temperature measurements. Due

to the exponential relationship between temperature and thermal dose, these artifactual

oscillations would result in overestimated thermal dose. When applied as input for closed-loop

temperature control, these fluctuations would also cause excessive fluctuations in output power,

amplifying artifactual variations in temperature measurement with real fluctuations in heating.

By averaging over a six-image sliding window, temperature standard deviation measured before

heating was reduced from 1.0°C to 0.4°C, matching the temperature measurement noise

observed without motion, but possibly reducing real temperature maxima and minima resulting

in an underestimation of thermal dose. Image artifacts due to transducer motion would be

completely eliminated by the use of electronic scanning with a stationary phased-array

transducer, such as those available with clinical focused ultrasound systems [255].

The use of rectal temperature measurements as the baseline temperature in MRI

thermometry may have added a constant offset to temperature measurements in muscle. A

combination of factors, including heat conduction, variations in water bath temperature, and

phase drift resulted in an observed temperature increase of approximately 1°C in untargeted

muscle regions.

In these preliminary experiments, closed-loop heating was allowed to continue in cases

where animal motion invalidated MRI temperature measurements, resulting in coagulation of

large regions centered on the target. In future work, these sudden movements could be

immediately detected in magnitude reconstructions of thermometry images to abort treatment

and prevent uncontrolled heating.

2.4.3 Hyperthermia mediated drug delivery

The results of this study demonstrate the feasibility of using MRI-guided focused ultrasound

hyperthermia with temperature-sensitive liposomes to achieve localized deposition of the

69

anticancer drug doxorubicin in a prescribed region of rabbit thigh. Doxorubicin concentrations

were significantly higher in heated than unheated tissues sampled from normal thigh muscle, and

in one rabbit, doxorubicin concentrations sampled at locations 15 mm or more away from the

boundary of the 10 mm diameter heated region were no different than in the unheated

contralateral thigh. These results were obtained using similar LTLD and thermal dose

prescriptions to those being used in Phase I/II clinical trials of LTLD combined with microwave

hyperthermia in the treatment of breast cancer recurrence at the chest wall.

In previous work using thermosensitive liposomes and regional rather than localized

hyperthermia in preclinical tumour models [68, 76, 181], passive drug carrier accumulation

augmented the localized enhancement in tumour drug concentration caused by triggered release.

These and other studies using NTLD in various rodent tumours showed that hyperthermia

increases both the rate and uniformity of liposome extravasation in angiogenic tumour vessels

[62, 84, 86, 129, 130, 135]. Hyperthermia is thought to cause morphological changes in tumour

endothelial cells, increasing the occurrence and size of pores in the vessel wall, enabling and

increasing extravasation of normally impermeable particles such as 100 nm liposomes [129,

130]. This increased extravasation of particles as large as liposomes is not seen in the vasculature

of healthy muscle tissue [129], suggesting that in the present study fluorescence measured in

tissue after the vessels were perfused with saline may have come from bioavailable doxorubicin

released from liposomes in the vasculature of the heated region. Without saline perfusion, intact

liposomes would continue to circulate with a plasma half-life of approximately 1 hour [188], and

the non-bioavailable doxorubicin trapped in these intravascular liposomes would increase the

fluorescence signal measured in bulk tissue samples at the 2 hour time point. Due to the small

fraction of the body being heated, free doxorubicin released in the vasculature of the target

region during heating is expected to have a minimal impact on systemic doxorubicin

concentrations, even compared to the small amount of doxorubicin that leaks from the liposomes

at body temperature [326]. Free doxorubicin in the vasculature is rapidly taken up by tissue with

an initial half-life of 8 minutes, followed by slow elimination with a terminal half-life of 30

hours [327].

70

2.4.4 Limitations and sources of variability in drug delivery

The LTLD formulation used in this study releases more than 80% of its payload in less than 20

seconds of exposure to temperatures greater than 41°C [302, 326], and clears from the

bloodstream with an initial half-life of about 1 hour [188], making release likely to occur as the

liposome passes through the heated region. In experiments where mild hyperthermia was

achieved, the duration over which the mean target temperature was greater than 41°C was 24 ± 5

minutes. In only two of these cases did mean temperatures ever exceed 45°C, and in those cases

it was for less than two minutes. In experiments where motion occurred, controlled temperatures

greater than 41°C were maintained for between 1 and 4 minutes prior to sudden motion resulting

in high temperatures due to controlled heating based on invalid temperature measurements. In

these cases, high power sonication continued for an average of 20 ± 6 minutes.

In cases where motion occurred and a high thermal dose was delivered, coagulative tissue

damage may have changed the mechanism for drug accumulation. In some cases, hyperthermia

may have allowed triggered release in healthy vessels prior to vascular shutdown. In other cases,

high temperatures experienced early on may have caused liposome extravasation due to vessel

damage, or rapid vessel collapse preventing liposomes from reaching the target tissue.

Coagulation could also have prevented the removal of residual liposomes from damaged vessels

by saline perfusion, exaggerating measurements of tissue drug concentration. These effects vary

spatially within a heated region, and the locations from which tissue samples were collected

created additional variability. In previous reports where non-thermosensitive liposomes were

combined with radiofrequency ablation in canine tumours, normal rat muscle, and normal rabbit

liver, high temperature ablation resulted in increased doxorubicin accumulation at the periphery

of the heated region compared to the central region of coagulative necrosis [147, 325]. The

efficacy of LTLD combined with radiofrequency ablation to treat hepatocellular carcinoma is

being investigated in a series of clinical trials [188].

For each experiment, total drug accumulation in heated and unheated regions was

reported in Table 2.1 as the average across 3-4 coarsely sampled 25-50 mg tissue sections.

Variability in measured drug concentration between samples from the same region may be

71

caused both by underlying variations in tissue drug concentration and sample loss during

analysis. These measurements do not account for the spatial distribution of doxorubicin released

in the microenvironment of the heated region, which requires further study using fluorescence

microscopy in normal tissue and tumour models. Previous tumour studies demonstrated

preferential liposome accumulation in some tumour regions, increased by hyperthermia [130],

but sequestered by their size to the perivascular space [129]. Rapid triggered release of large

concentrations of drug in the tumour vasculature and perivascular space may lead to the

subsequent diffusion of free doxorubicin into deeper cell layers for greater antitumour effect.

Further studies to characterize the spatial distributions of liposomes and bioavailable drug in the

tumour microenvironment are required to evaluate the benefits of using MRI-guided FUS

hyperthermia to mediate localized drug release.

2.5 Conclusions Temporally and spatially uniform hyperthermia was localized to a target region in rabbit thigh

muscle using mechanically-scanned focused ultrasound under multi-point proportional-integral

MRI temperature control. Administration of lyso-thermosensitive liposomal doxorubicin during

localized hyperthermia resulted in localized drug release. Measured doxorubicin concentrations

were significantly higher in muscle tissue sampled from heated regions two hours after exposure.

The results demonstrate the potential of MRI-controlled focused ultrasound hyperthermia for

enhanced local drug delivery with temperature-sensitive drug carriers. Further experiments are

required to determine the effects of temperature, heating duration, and thermal dose on drug

deposition in heated regions, and on the distribution of released drug in the tumour

microenvironment.

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Chapter 3 MRI-controlled focused ultrasound hyperthermia for targeted drug delivery in bone: in vivo results2

3.1 Introduction Bone metastases occur in approximately 70% of patients with advanced breast and prostate

cancer [328]. Their sequelae impair a patient’s quality of life with pathologic fractures, spinal

cord compression, hypercalcemia, and debilitating pain. For localized pain related to skeletal

metastases, palliative radiotherapy is the standard of care, but 20-30% of patients do not

experience adequate relief [329]. For some patients, bisphosphonate therapy has shown promise

in breaking the cycle of tumour progression and osteoclast-mediated bone resorption, decreasing

and delaying the incidence of skeletal complications and pain [330, 331].

In patients with intractable pain refractory to radiotherapy and bisphosphonates or

systemic chemotherapy, thermal ablation of osteolytic bone metastases using percutaneous

radiofrequency ablation (RFA) under ultrasound or computed tomography image-guidance

achieves pain relief by ablating periosteal nerve endings and reducing tumour burden [332, 333].

Recently, thermal ablation using MRI-guided focused ultrasound demonstrated effective

palliation of pain arising from both osteolytic and osteoblastic bone metastases in patients who

had exhausted all other treatment options [238, 239, 334]. Focused ultrasound has several

advantages over RFA, most important of which is noninvasive heating with the ability to cover

larger lesions.

2 This chapter is adapted from the article: “Hyperthermia in bone generated with MR imaging–controlled focused ultrasound: control strategies and drug delivery”. Robert Staruch, Rajiv Chopra, Kullervo Hynynen. Radiology 263(1): 117-127 (2012).

74

In patients with painful bone metastases, combining localized drug release with thermal

pain palliation could increase cell kill at treatment margins and decrease the required energy to

achieve a therapeutic effect, potentially reducing treatment time and risk of normal tissue

damage. Localized drug release mediated by hyperthermia, rather than high-temperature thermal

ablation, may reduce pain related to treatment and possibly achieve local tumour control without

compromising skeletal stability. This noninvasive localized therapy would depend on the ability

to maintain desired temperatures in targeted regions, with minimal heating of surrounding tissue.

Focused ultrasound under closed-loop feedback control based on MR thermometry has

been used to achieve precise noninvasive heating in soft tissue [289, 290, 308], and MRI has

been used to target and evaluate the effects of focused ultrasound on bone [335, 336].

Temperature control in bone is made difficult by the inability of proton resonance frequency

(PRF)-shift MR thermometry to measure temperature in cortical bone. In this study, a feedback

control system is proposed that incorporates PRF thermometry to maintain desired temperatures

in bone based on temperatures measured in adjacent soft tissue. Its ability to generate localized

hyperthermia in bone for localized drug delivery was evaluated. Two different control scenarios

were investigated, demonstrating the effect of increased ultrasound absorption in bone on

temperature control and drug delivery.

The purpose of this study was to evaluate the feasibility of achieving image-guided drug

delivery in bone using MRI-controlled focused ultrasound hyperthermia and temperature-

sensitive liposomes.


3.2.1 Animals

Nine male New Zealand white rabbits weighing 3-4 kg were anaesthetized by intramuscular

injection of ketamine (50 mg/kg/hr) and xylazine (10 mg/kg/hr). Anaesthetized rabbits had one

ear vein cannulated, both thighs depilated, and were placed on a stage above the degassed water

tank of an MRI-compatible focused ultrasound system (Figure 3.1). During hyperthermia, rectal

75

temperature was monitored by a fiber-optic temperature probe (3100, Luxtron Corp., Santa

Clara, CA), and maintained by manually regulating the temperatures of a hot water blanket

covering the animal and the degassed water reservoir below.

3.2.2 Focused ultrasound system

Acoustic energy was delivered by continuous sonication with a spherically-focused, air-backed

ultrasound transducer (fundamental frequency, 0.536 MHz; curvature radius, 10 cm; aperture

diameter, 5 cm) driven at its fifth harmonic (2.787 MHz) as described in Chapter 2.

The transducer was incorporated into an MRI-compatible positioning system similar in

function to that described in [313], programmed to move along a 10 mm diameter circular

trajectory at a speed of 1 revolution per second, for simultaneous ultrasound heating and MR

thermometry. Degassed water coupled the transducer through a window of 25 μm polyimide film

(Kapton®, DuPont, Wilmington, DE) into the target thigh. The thigh was positioned such that

the ultrasound beam penetrated the skin at approximately normal incidence and was focused at

the muscle-bone interface. To prevent undesired tissue heating in the contralateral thigh, a saline

bag and polyurethane rubber acoustic absorber (AptFlex F28, Precision Acoustics, Dorchester,

UK) were placed between the rabbit’s thighs [287].


Experiments were performed with the positioning system placed in a clinical 1.5T MRI (Signa,

GE Healthcare, USA). MRI-controlled focused ultrasound was used to heat 10 mm diameter

regions of thigh muscle near the muscle-bone interface to 43°C for 20 minutes. This temperature

elevation and exposure duration were chosen to be sufficient to trigger drug release from

temperature-sensitive liposomes [301] without causing significant tissue damage [106].

A single-loop receive coil with a square opening of 85 mm was placed underneath the

animal around the polyimide film window to improve the signal to noise ratio (SNR) in the

heated region. Before heating, anatomical images were acquired with T1-weighting and T2-

weighting (Table 3.1), for target definition and verification of acoustic coupling. These images

76

Figure 3.1: Experimental setup for MRI-controlled focused ultrasound heating in bone.A) Axial localizer image along the ultrasound beam path, indicating the location of thethree coronal planes used for MR thermometry, two of which are shown in B) and C). B)Coronal fast spin-echo T2-weighted image (2000/75 [repetition time ms/echo time ms]) inmuscle near the bone interface, in which MR thermometry was used to control treatment.C) Coronal T2-weighted image through the bone, in which MRI thermometry was used todetect excessive heating adjacent to the bone.

77

Table 3.1: MR imaging parameters.

Sequence*

Repetition Time (ms)

Echo Time (ms)

Flip Angle (deg)

Echo Train

Length NEX

Field of

View (cm)

Matrix Size

Section Thickness

(mm) Bandwidth

(kHz) Imaging Plane(s)

FSE-T2 2000 75 90 4 2 16 256x128 3 15.63 Coronal, Axial

FSE-T1 500 15 90 2 3 16 256x128 3 15.63 Coronal, Axial

FSPGR 38.6 10 30 N/A 1 16 128x128 5 31.25 Coronal

*All MR imaging examinations were performed at 1.5 T (Signa, GE Healthcare, USA). T2 and T1-weighted fast spin-echo (FSE) images were acquired for treatment planning, with no phase wrap. Fast spoiled gradient-echo (FSPGR) images were acquired for thermometry.

were re-acquired after heating to evaluate changes in tissue perfusion and/or tissue damage

caused by ultrasound heating, identified as regions of either increased signal intensity or

decreased signal intensity with a gadolinium-enhanced rim. Post-sonication T1-weighted images

were acquired before and 1 minute after bolus injection of MRI contrast agent in the ear vein (0.2

mmol/kg gadodiamide, Omniscan; GE Healthcare).

During mechanical scanning of the ultrasound transducer, coronal RF-spoiled gradient-

echo images were continuously acquired in 3 planes: at the location of the bone, in the muscle

directly beneath the bone, and in the muscle closer to the skin (Figure 3.1). When new images

were available (every 5 seconds), a real-time acquisition interface [317, 318] transferred k-space

data to a control computer for image reconstruction and calculation of spatial temperature

distribution using software written in MATLAB (Mathworks, Natick, MA). Phase difference

maps were calculated by complex phase subtraction between treatment images [268] and the

average of five baseline images acquired during transducer motion prior to heating. Temperature

maps were obtained from phase difference maps using a PRF-shift coefficient of -0.010 ppm/°C

[268] and adding the baseline temperature measured by a fiber-optic probe in the rectum.

Temperatures averaged over a 25-35 mm2 image region in a mineral oil reference phantom were

subtracted to account for magnetic field drift during the experiment [268]. Temperature images

were then averaged over a 30 second sliding window, which was shown previously to be

78

effective in reducing the effect of periodic susceptibility-related phase shifts caused by circular

transducer motion during imaging [337].

Spatial temperature elevations in muscle just below the bone interface were controlled by

eight independent proportional-integral feedback controllers along the scanned heating

trajectory, rapidly switching applied power as the transducer scanned along the circular

trajectory, and updating controller outputs after each image acquisition according to the

following equation:

[ ] [ ]( ) [ ]( )∑=

−+−=n

igoalIgoalP iTTKnTTKnP

18..18..18..1 . [Eq. 7]

For temperature image n, the difference between the temperatures T1..8 in the eight pre-defined 1

mm2 control regions and the target temperature Tgoal, as well as the integral of the difference over

each previous image i, were scaled by proportional and integral gains KP and KI to specify the

acoustic powers P1..8 delivered by the transducer as the focus crossed each region. The average

lag time for data transfer, reconstruction, and controller update was measured to be less than one

second.

Two different control scenarios were investigated. In the first, the ultrasound focus and

the image plane used for temperature control were set in muscle 10 mm below the bone interface

(10 mm offset), using gain settings reported previously for controlled heating in muscle (KP =

4.5, KI = 0.03, [337]). This case emulates the scenario where there is bone located behind a

targeted tumour. In the second scenario, the control plane was defined immediately under the

bone with the focus centered at the bone interface (0 mm offset); gains were selected using

simulations of scanned focused ultrasound heating (KP = 3.0, KI = 0.0). For each strategy,

simulations were also used to estimate temperature elevations in bone and marrow.

79

3.2.4 Simulations of temperature elevations in bone

Three-dimensional numerical simulations were used to estimate temperature elevations in bone

and to identify appropriate controller gain values for scanned focused ultrasound heating near the

muscle-bone interface. The absorbed acoustic power was calculated using a multilayer

transmission model of the acoustic field pattern [321, 322] as shown in Figure 3.5A and the

tissue properties listed in Table 3.2. As in the in vivo experiments, the acoustic focus was set

either at the muscle-bone interface or at an axial offset of 10 mm below the bone, shifted off-

center by the 5 mm scan radius and moved in a circular path around the center using affine

transformations and linear interpolation. These two locations were chosen to determine the

optimal depth for MRI-controlled hyperthermia that achieved uniform hyperthermic

temperatures in the soft tissue, and controller gain values were optimized for each scenario. At

each 0.05 second time step, the three-dimensional temperature distribution was recalculated by

the Pennes bioheat transfer equation [214] using a previously described finite-difference time-

domain implementation [322], with a grid size of 0.25 x 0.25 x 1.0 mm and boundary conditions

of water at 37°C on all surfaces. To model feedback control, temperatures on the simulation grid

were averaged spatially and temporally, and zero-mean Gaussian noise with a standard deviation

of 0.5 was added to approximate the spatial resolution, sampling rate and temperature

measurement noise of MR thermometry (1 x 1 x 5 mm, 5 seconds, ±0.5°C, respectively).

3.2.5 Drug administration and analysis of drug concentrations in tissue

Lyso-thermosensitive liposomal doxorubicin (LTLD, Thermodox®, Celsion Corporation,

Columbia, MD) was provided by the manufacturer. Rabbits were administered a doxorubicin

dose of 2.5 mg/kg diluted as equal parts LTLD and 5% dextrose sterile solution, infused at 1.2

ml/min into the ear vein during hyperthermia, starting when the average temperature in the target

region reached 43°C. One leg was heated in each rabbit; the other served as the control. In four

rabbits, the focus and the image plane used for control were offset 10 mm from the femur; in five

rabbits, a 0 mm offset was used.

80

Table 3.2: Acoustic and thermal properties for numerical simulations of scanned focused ultrasound heating in bone.

Material property Water Muscle Cortical bone

Trabecular bone

Thickness (mm) 80 or 90 20 and 40 2 6

Density ρ (kg m-3) 1000 1060 1700 1300

Speed of sound c (m s-1) 1500 1572 1333 1500

Specific heat capacity ct (J kg-1 °C-1) 4180 3720 1600 1600

Thermal conductivity k (W m-1 °C-1) 0.615 0.537 0.6 0.6

Attenuation αatt (Np m-1 MHz-1) 2.88 x 10-4 4.1 250 5.8

Perfusion wb (kg m-3 s-1) 0.0 1.0 0.0 0.5

Two hours after LTLD infusion, unabsorbed liposomes were flushed from the vasculature

by transcardiac saline perfusion before collection of tissue samples for evaluation of drug

release. After sacrifice, 10 mm thick axial sections of the femur with its marrow, as well as 50-

100 mg samples of adjacent muscle, were harvested from the targeted region of the treated leg

and matching regions of the untreated leg, snap-frozen in liquid nitrogen, and stored at -80°C.

Tissue doxorubicin concentrations were measured by the fluorescence intensity of

doxorubicin extracted from homogenized tissue samples, using a technique based on that

described in Chapter 2 [58]. Tissue samples were weighed and added to 20 volumes of acidified

ethanol extraction solvent (0.3N HCl in 50% ethanol) before homogenization with a tissue

grinder (PYREX® Ten Broeck, Corning, USA), overnight refrigeration in acidified ethanol

solvent, centrifugation (16000g, 30 min), and storage of supernatants in the dark at -20°C. 1.5 ml

of acidified ethanol was added to 0.5 ml of supernatant in 3 ml fluorometry cuvettes, and

fluorescence intensity was measured using a benchtop fluorometer (VersaFluor, Bio-Rad

Laboratories, Hercules, CA) with 480 nm excitation and 590 nm emission filters. Relative

81

fluorescence intensities were scaled to doxorubicin concentrations using a fluorescence

calibration curve of a serial dilution of free doxorubicin (Doxorubicin HCl, Teva Novopharm,

Toronto, ON) added to 0.5 ml blank tissue homogenates in 1.5 ml of acidified ethanol.

3.2.6 Statistical Analysis

Descriptive statistics were calculated using means and standard deviations. The one-sided

Wilcoxon signed-rank test was used to identify drug concentration enhancements in heated vs.

unheated tissues taken on the same rabbit for samples of marrow and muscle at the bone

interface (GraphPad Prism 5.0; GraphPad Software, La Jolla, CA). Differences were considered

statistically significant for values of p < 0.05.

3.3 Results


Figures 3.2A and 3.2B show temperature distributions in the same coronal imaging planes

depicted in Figures 3.1B and 3.1C, with temperature greater than 37°C overlaid on the

corresponding magnitude image, at a selected time of 10 minutes after heating began. In Figure

3.2B, pixels in bone where SNR was too low to produce accurate temperature measurements

were masked out for display. In this experiment, the focus was set at the muscle-bone interface

(0 mm offset), with the top of the 5 mm thick imaging plane used for temperature control set just

below the bone. Temperature evolution within the 10 mm targeted region of the control plane is

shown in Figure 3.3, as well as the power applied at each control point over time. At each time

point, the mean, T90 and T10 temperatures are shown, as well as the mean temperatures in

unheated regions 10 to 12.5 mm from the edge of the circular target, the controller input function

Tgoal, and the start/end times of LTLD infusion. Figure 3.4 shows the radial mean ± SD of the

steady-state temperature distribution and median thermal dose, centered on the target region, for

the control plane and the bone. Hyperthermia experiments are summarized in Table 3.3, showing

average steady-state mean, T90 and T10 temperatures of 43.2°C, 42.0°C, and 44.4°C, with

standard deviations of 0.4°C over the experiment duration.

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Figure 3.2: Snapshot of temperature distribution from controlled hyperthermia at muscle-bone interface in rabbit thigh using MRI-controlled focused ultrasound. Control regions of interest, circular unheated region of interest, and temperatures greater than 37°C are overlaid on coronal FSPGR magnitude images (38.6/10, 30° flip angle) acquired A) in muscle used for feedback control, and B) in bone, with low signal bone pixels masked. Images acquired 10 min after start of heating.

3.3.2 Numerical simulations of temperature elevations in bone during

MRI-controlled hyperthermia

Temperature elevations in bone caused by scanned focused ultrasound heating with beam offsets

of 0 and 10 mm from the bone were estimated using numerical simulations over the tissue

volume depicted in Figure 3.5A. Calculated temperatures were averaged over the image planes

shown. Figure 3.5(B, top) shows coronal temperature distributions in the control plane for the 0

mm and 10 mm offset scenarios, both with average control region temperatures near 43°C,

similar to in vivo results. However, the respective sagittal images in Figure 3.5(B, bottom) show

higher temperatures in bone and adjacent soft tissues for the 10 mm offset. In this situation,

feedback control is based only on temperatures measured in the focal plane, maintaining the

desired temperature of 43°C in the soft tissue of the target region 10 mm from the bone.

However, undesired thermal damage occurs at the bone surface due to increased ultrasound

absorption in bone, with heat spreading away from the bone by conduction over the 20 minute

heating duration. The median, T90 and T10 temperatures within 10 mm diameter regions centered

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Figure 3.3: Temporal evolution of controlled hyperthermia at muscle-bone interface in rabbit thigh using MRI-controlled focused ultrasound, for the same experiment shown in Figure 3.2. Top: Mean, T90 and T10 temperatures measured in the 10 mm diameter target region in the plane used for temperature control. Liposome infusion duration is shaded. Bottom: Acoustic power applied at each of the eight control regions at each time step.

Figure 3.4: Radial distribution of thermal dose and temporal mean ± SD steady-state temperature in A) control plane and B) bone plane, for the same experiment shown in Figures 3.2 and 3.3. Bone pixels are masked and the unheated region of interest is shaded.

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Table 3.3: Summary of in vivo MRI-controlled focused ultrasound hyperthermia experiments with the focus set at two different offsets from a muscle-bone interface in rabbit thigh.

Rabbit Target Radius

Mean ± SD temperature

Mean ± SD T90

Mean ± SD T10

Radius at 42°C

(muscle)

Radius at 41°C

(bone)

Thermal dose

(muscle)

Thermal dose

(bone)

(mm) (°C) (°C) (°C) (mm) (mm) (CEM43) (CEM43)

Control plane 10-15 mm away from bone interface (n = 4)

1 5.0 43.3 ± 0.4 42.5 ± 0.4 44.1 ± 0.4 5.9 12.0 28.0 112.0

2 5.0 43.2 ± 0.3 42.3 ± 0.3 44.0 ± 0.3 5.7 6.6 24.2 12.8

3 5.0 43.2 ± 0.3 42.2 ± 0.3 44.0 ± 0.4 5.3 7.0 28.0 126.3

4 5.0 43.7 ± 0.6 42.4 ± 0.6 45.0 ± 0.7 5.5 4.4 45.1 5.8

Mean ± SD 43.3 ± 0.3 42.3 ± 0.4 44.3 ± 0.5 5.6 ± 0.3 7.6 ± 3.3 31.3 ± 9.4 64.2 ± 63.7

Control plane at bone interface (n = 5)

5 5.0 43.1 ± 0.3 41.8 ± 0.4 44.4 ± 0.4 5.3 4.7 29.8 8.4

6 5.0 42.7 ± 0.3 41.2 ± 0.3 44.1 ± 0.3 5.3 2.9 20.4 1.3

7 5.0 43.2 ± 0.5 41.9 ± 0.5 44.6 ± 0.6 4.7 4.7 27.8 8.9

8 5.0 43.3 ± 0.3 41.5 ± 0.3 44.8 ± 0.3 4.9 5.5 32.9 5.6

9 5.0 43.0 ± 0.3 41.8 ± 0.4 44.1 ± 0.4 5.1 5.2 24.5 11.8

Mean ± SD 43.1 ± 0.3 41.6 ± 0.4 44.4 ± 0.4 5.0 ± 0.3 4.5 ± 1.2 26.4 ± 5.3 6.9 ± 4.6 Temperature (mean across 10 mm diameter target region), T90 (temperature which 90% of target region exceeds), and T10 (temperature which 10% of target region exceeds) are reported as the average over all treatment temperature images starting once the target temperature reached 95% of the desired temperature elevation. CEM43 = cumulative equivalent minutes at 43°C.

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around the beam axis in cortical bone, marrow, and 5 mm thick regions of muscle centered at

depths of 2.5 mm and 12.5 mm below the bone are summarized in Table 3.4, showing bone

temperature elevations 4.2 to 4.7 times higher than the control region for the 10 mm offset case,

as opposed to 1.1 to 1.8 times higher for a 0 mm offset. These elevated temperatures in the 10

mm offset case resulted in thermal damage identified on post-treatment T2 and contrast-

enhanced T1-weighted imaging. Temperature variation in cortical bone results from high

temperatures along the scan trajectory where applied power is dynamically modulated versus the

central regions heated by conduction, as shown in simulated radial temperature distributions in

Figure 3.6.

Figure 3.5: Numerical simulations of MRI-controlled focused ultrasound hyperthermia at a muscle-bone interface. A) Simulation volume including layers of water, muscle, cortical bone, and trabecular bone, using tissue properties listed in Table 3.2. B) Left: Average steady-state temperatures for experiment where focal plane is offset 0 mm from bone interface, calculated in control plane 0 mm from interface (top) and sagittal plane (bottom). B) Right: Average steady-state temperatures for experiment where focal plane is offset 10 mm from bone interface, calculated in control plane 10 mm from interface and sagittal plane.

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Figure 3.6: Numerically simulated average steady-state temperature vs. radius for two control scenarios: control plane set 0 mm from bone interface (empty markers), and control plane offset 10 mm from the bone interface (filled). For each scenario, average steady-state temperatures were calculated in cortical bone, marrow, muscle adjacent to the bone, and muscle 10 mm from the bone interface. For clarity, every fourth marker is shown.

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Table 3.4: Numerical simulations of MRI-controlled focused ultrasound hyperthermia in bone, with the focus set at two different offsets from the muscle-bone interface.

Simulation Plane Mean ± SD T50

Mean ± SD T90

Mean ± SD T10

Radius at 42°C

Radius at 41°C

Thermal dose

(°C) (°C) (°C) (mm) (mm) (CEM43)

Control plane offset 10 mm from bone interface

Muscle 10mm offset 43.1 ± 0.2 43.0 ± 0.2 43.2 ± 0.2 5.6 6.6 23.2

Muscle 0mm offset 66.1 ± 0.2 64.5 ± 0.3 66.6 ± 0.3 13.6 14.4 2.6 x 108

Cortical Bone 78.1 ± 0.8 75.8 ± 0.6 80.1 ± 1.1 14.3 15.1 6.6 x 1012

Marrow 58.0 ± 0.3 56.5 ± 0.3 59.3 ± 0.3 13.5 14.5 8.0 x 105

Control plane offset 0 mm from bone interface

Muscle 10mm offset 36.0 ± 0.2 35.8 ± 0.2 36.1 ± 0.2 0 0 0.0

Muscle 0mm offset 43.2 ± 0.6 42.5 ± 0.6 43.9 ± 0.7 5.4 6.2 28.7

Cortical Bone 46.8 ± 2.0 44.9 ± 1.4 50.2 ± 3.5 7.2 7.8 2.0 x 103

Marrow 40.7 ± 0.4 40.2 ± 0.5 41.1 ± 0.4 0 4.0 1.2

3.3.3 Doxorubicin concentrations in bone marrow and muscle

Table 3.5 shows the doxorubicin concentrations measured in tissue samples from each rabbit,

collected from the marrow and adjacent muscle in both the targeted region of the treated thigh

and equivalent regions in the untreated contralateral thigh. Overall, the fold-increases of

doxorubicin concentration in heated vs. unheated marrow and muscle were 8.2 ± 3.7 times (mean

± SD) and 16.8 ± 6.9 times, respectively. Figure 3.7A shows the mean ± SD doxorubicin

concentrations for each tissue type across all experiments; paired differences were statistically

significant for each tissue type (marrow, p = 0.002; muscle, p = 0.002, one-sided Wilcoxon

signed-rank test).

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Figure 3.7B divides doxorubicin concentration data into experiments with the control

plane offset 10-15 mm from the bone (causing uncontrolled heating and thermal coagulation at

the bone interface), versus 0 mm from the bone (controlled hyperthermia of the tissues

surrounding the muscle-bone interface). With the control plane at the bone interface, increases of

doxorubicin concentrations in heated vs. unheated marrow and muscle were 9.9 ± 3.3 times (p =

0.03) and 18.2 ± 8.7 times (p = 0.03), respectively. Enhancements were seen when the control

plane was set 10-15 mm from the bone interface but did not reach statistical significance, with

increases of 6.1 ± 3.3 times (p = 0.06) for marrow and 15.1 ± 4.1 times (p = 0.06) for muscle.

Figure 3.7: Doxorubicin concentrations measured by fluorescence intensity in tissuesamples harvested from heated and unheated regions of rabbit thigh muscle and bonemarrow. A) Averaged across all experiments, and B) separated into experiments wherefocal offsets of 10 mm and 0 mm were used. Columns, mean; bars, SD.

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Table 3.5: Doxorubicin concentrations measured by fluorescence intensity in tissue samples harvested from heated and unheated regions of rabbit thigh muscle and bone marrow.

Rabbit DOX in Muscle [ng/mg] DOX in Marrow [ng/mg]

Unheated Heated SS Unheated Heated SS

Control plane 10-15 mm away from bone interface (n = 4)

1 1.6 19.3 8.5 29.6

2 2.2 24.5 5.9 28.4

3 2.1 40.3 4.2 45.5

4 2.0 36.0 8.3 42.4

Mean ± SD 2.0 ± 0.3 30.0 ± 9.8 p=0.06 6.7 ± 2.1 36.5 ± 8.8 p=0.06

Control plane at bone interface (n = 5)

5 1.7 21.9 4.3 41.6

6 1.6 33.2 4.4 22.1

7 2.2 40.6 5.3 56.9

8 3.5 28.1 4.0 39.0

9 1.2 37.3 3.8 54.1

Mean ± SD 2.0 ± 0.9 32.2 ± 7.4 p=0.03 4.3 ± 0.6 42.7 ± 13.9 p=0.03

Overall (n = 9)

Mean ± SD 2.0 ± 0.7 31.2 ± 8.0 p=0.002 5.4 ± 1.8 40.0 ± 11.7 p=0.002 One-sided Wilcoxon signed-rank test, p < 0.05 considered statistically significant (SS). DOX = doxorubicin.

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3.4 Discussion

3.4.1 Hyperthermia and drug delivery in bone

In our study, image-guided localized drug delivery in bone was achieved noninvasively using

MRI-controlled focused ultrasound heating and temperature-sensitive liposomes. Temperatures

of 43°C were generated in a 10 mm diameter circular region at a bone interface using scanned

focused ultrasound, and maintained for 20 minutes by controlling ultrasound power based on

MRI temperature measurements in adjacent soft tissue. Administration of lyso-thermosensitive

liposomal doxorubicin during heating resulted in locally-enhanced drug deposition in bone

marrow and muscle adjacent to the bone surface.

The proposed method controls localized bone heating in an automatic feedback-control

loop using PRF-shift MR thermometry in adjacent soft tissue. The PRF-shift technique is based

on temperature-dependent changes in the local magnetic field around water protons, and is of

limited use in tissues with low water content such as bone, fat, and lung. Phase subtraction is also

sensitive to tissue displacements in the image plane and magnetic field distortions related to

tissue motion near the image plane, limiting the proposed method to targets immediately adjacent

to non-moving aqueous tissue where stable thermometry can be achieved. Sprinkhuizen et al

[338] have demonstrated that temperature-dependent magnetic susceptibility changes in fat

(0.0094 ppm/°C) influence the PRF-shift field changes in nearby water; we calculate that the

temperature-dependence of bone’s susceptibility is much smaller (0.0007 ppm/°C for a

susceptibility of -8.86 ppm [319] and linear thermal expansion coefficient of -0.27 x 10-4/°C

[339]), and will not affect the field of neighbouring soft tissue during heating. With externally-

focused ultrasound heating, this method is limited to targets with an acoustic window through

soft tissue. At high frequencies (3 MHz), ultrasound-generated temperature elevations occur

close to the bone interface, and temperatures in cortical bone can be approximated by

temperatures measured in adjacent soft tissue. This was confirmed by our simulations at 2.787

MHz; simulated temperatures in muscle also demonstrated good agreement with in vivo MR

temperature measurements. Simulations using a higher speed of sound in bone would increase

reflected power at the bone interface, resulting in a smaller difference between temperatures in

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muscle and bone, and lower temperatures in the marrow and distal bone. At lower frequencies (1

MHz), peak temperatures in bone will exceed adjacent soft tissue and occur further from the

bone interface [215, 340], which may be useful when heating thick layers of cortical bone in

osteoblastic lesions. MRI-controlled bone heating with other energy sources, including invasive

radiofrequency applicators, would be limited by interference between electromagnetic heating

devices and MR imaging, and by poor electrical conductivity (and thus heating) in bone [192].

Previous preclinical studies using thermosensitive liposomes have demonstrated 10 to 20-

fold increases in drug concentration in tissue samples harvested from animal tumours and muscle

heated by water [76, 177], microwave applicators [181], and ultrasound [178, 337]. With non-

thermosensitive liposomes, similar increases have also been seen in the periphery of lesions

created in tumours, kidney, and liver using radiofrequency ablation [147]. The results of our

study demonstrate the feasibility of achieving image-guided drug deposition near bone,

achieving 6 and 15-fold increases in heated vs. unheated bone marrow and muscle adjacent to

bone, respectively. In our study liposomes were administered during heating and drug deposition

was quantified using the fluorescence intensity of released doxorubicin in homogenized bone

marrow and muscle samples. Clinical pharmacokinetic data suggests that greater drug deposition

may be possible if heating is administered 30-60 minutes after infusion [13]. In cases of thermal

damage, vascular shutdown presumably acted both to prevent liposomes from reaching the target

region and to prevent released drug in tissue from returning to the vasculature, again highlighting

the importance of infusion timing and target temperature. Future studies will focus on

optimization of infusion and heating protocols, and on performing histology to determine the

spatial distribution of doxorubicin with respect to patterns of thermal damage in heated tissue.

3.4.2 Practical Applications

In clinical studies of MRI-guided focused ultrasound thermal ablation in bone, MR thermometry

has been used to monitor soft tissue temperature elevations during treatment to enable the

clinician to halt or adjust treatment parameters and avoid unintended thermal damage to

surrounding tissue [238, 239, 334, 336]. Commercial systems are capable of closed-loop MR

thermometry control in soft tissue [280], and could implement the proposed method directly for

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controlled heating in bone. Our results suggest that with the focus set at the bone interface, bone

temperatures in clinical treatments could be safely controlled using temperatures measured in

adjacent soft tissue. They also demonstrate the importance of multi-plane thermometry to

observe heating outside the ultrasound focus, as already implemented commercially [255].

Clinical trials are currently underway combining thermosensitive liposomal doxorubicin

with both thermal ablation using percutaneous RFA for hepatocellular carcinoma [190], and with

mild heating using microwave hyperthermia for recurrent breast cancer at the chest wall [185]. In

our study, tissue drug concentrations were increased in heated muscle and marrow both for

experiments where the focus was set 10 mm away from the bone and large regions of thermal

damage were observed, and where the focus was set 0 mm from the bone and thermal damage

was limited to a narrow region at the bone interface. This suggests that clinical treatments could

use mild heating to avoid thermal damage near critical structures while achieving a similar

degree of drug deposition as would be achieved with ablative heating. Further studies are

required to optimize infusion protocols and heating durations for drug deposition in large targets

that cannot be covered in a single sonication, and to determine whether local drug delivery to

bone without thermal coagulation of the soft tissue could achieve local pain relief or tumour

control.

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Chapter 4 Enhanced drug delivery in rabbit VX2 tumours using thermosensitive liposomes and MRI-controlled focused ultrasound hyperthermia3

4.1 Introduction The clinical efficacy of chemotherapy in solid tumours is impaired by systemic toxicity and the

inability of anticancer drugs to reach all cancer cells in sufficient concentration to cause

cytotoxicity [30]. To treat tumour cells that are inadequately perfused by their disorganized

vasculature, drugs must cross the vessel wall and penetrate the dense extracellular matrix,

overcoming sequestration in perivascular cells [31]. Anticancer drugs must reach tumour cells in

a bioavailable form at a cytotoxic dose; however, the administered dose must not exceed levels

that cause systemic toxicity arising from efficient delivery through highly organized vasculature

in well-perfused normal tissues such as the liver and kidney [24].

Encapsulating chemotherapeutic agents in long-circulating liposomal drug carriers

prevents extravasation from normal microvasculature, while allowing preferential extravasation

from leaky tumour vessels [129]. However, passive liposome accumulation occurs over a period

of 24-48 hours; to prevent toxicity from systemic drug release, liposomes are designed to

degrade slowly after accumulating in tumours, thereby delivering sustained, but low,

concentrations of the bioavailable drug [136]. Peak concentrations of the bioavailable drug at the

target site can be increased with thermosensitive liposomes, which clear from the bloodstream

with an initial half-life of about 1 hour, but release encapsulated cytotoxic agents when heated to 3 This chapter is adapted from the article: “Enhanced drug delivery in rabbit VX2 tumours using thermosensitive liposomes and MRI-controlled focused ultrasound hyperthermia”. Robert M. Staruch, Milan Ganguly, Ian F. Tannock, Kullervo Hynynen, Rajiv Chopra. International Journal of Hyperthermia 28(8): 776-787 (2012).

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a critical temperature of approximately 41°C [179, 180, 184, 301]. In preclinical studies using

regional heating in small animal tumour models, heat-triggered release results in increased

overall drug concentrations, reduced blood flow [183], and enhanced antitumour effect [177];

however, the underlying mechanisms remain unclear. One theory is that rapid drug release from

thermosensitive liposomes in heated tumour vessels creates a localized concentration gradient

from the vessel into the tumour interstitium, thereby increasing the penetration of anticancer

drugs through the tumour microenvironment [76]. With prolonged heating, thermally-triggered

release from intact liposomes continually circulating into the heated region could maintain

exposure to high levels of the bioavailable drug [326]. However, little is known about how

triggered release affects the microregional distribution of anticancer drugs in solid tumours in

vivo.

Previously, hyperthermia-mediated drug delivery using focused ultrasound has been

demonstrated in normal tissue [337], small animal tumour models [68, 179], and one recent

study demonstrating feasibility in rabbit tumours [341]. Here, we report the effects of thermally-

triggered doxorubicin release in rabbit tumours using a commercially-developed thermosensitive

drug carrier and MRI-controlled focused ultrasound hyperthermia. We demonstrate temperature

control across a large area of each tumour, compare drug concentrations in heated and unheated

tumours, and investigate the microregional distribution of released drug within the tumour

microenvironment.


4.2.1 Animals and VX2 tumours

Thirteen days before treatment, male New Zealand White rabbits (2.5-3.5 kg) were injected in

both thigh muscles with a 0.4 ml suspension of 3 to 4 million VX2 carcinoma cells in Hank’s

Balanced Salt Solution. One tumour was heated in each rabbit, the other served as the control.

Prior to ultrasound heating, rabbits were anaesthetized by intramuscular injection of

ketamine (50 mg/kg/hr) and xylazine (10 mg/kg/hr), and had one ear vein and one ear artery

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cannulated. Both legs were depilated to enable transmission of ultrasound into the thigh, and then

rabbits were placed on a stage above the degassed water tank of an MRI-compatible focused

ultrasound system (Figure 4.1A). During treatment, rectal and skin temperatures were monitored

by fiber-optic temperature probes (3100, Luxtron Corp., Santa Clara, CA), and maintained at the

same temperature by manually regulating the temperatures of a hot water blanket covering the

animal and the degassed water reservoir below (T/Pump Model TP-500, Gaymar, Orchard Park,

NY).


MRI-controlled focused ultrasound was used to heat 10 mm diameter tumour regions to 43°C for

20 minutes, using the methods described in Chapter 2. This temperature elevation and exposure

duration were chosen to be sufficient to trigger drug release from thermosensitive liposomes

[301] and increase local perfusion within the heated region of tissue [121], without causing

significant tissue damage [106].

Before treatment, anatomical MR images were acquired with T1-weighting and T2-

weighting to identify the tumour location, verify acoustic coupling, and to define a 10 mm

diameter target encompassing all or most of the tumour to be heated. The ultrasound transducer

was positioned to set the focus at the depth of the targeted tumour, and three imaging planes for

MR thermometry were prescribed perpendicular to the ultrasound beam, centered at the depth of

the focus. During heating, temperatures in the MR thermometry plane through the center of the

tumour were used for feedback control as the focus was scanned along the perimeter of the target

region, heating the interior of the region by conduction (Figure 4.1B) [337]. Following

sonication, T2-weighted and contrast-enhanced T1-weighted images (0.2 mmol/kg gadodiamide,

GE Healthcare) were acquired to evaluate tissue and perfusion changes related to thermal

damage and drug release.

The temporal and spatial uniformity of heating was evaluated based on temperatures

measured within MR thermometry image regions matching the intended target diameter, for all

images after the target temperature reached 42°C. The temporal average of the spatial mean, 10th

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percentile and 90th percentile temperatures at each time point are reported, as well as the

diameter receiving a time-averaged temperature of greater than 41°C, and the duration over

which a spatially-averaged temperature of greater than 41°C was achieved. Thermal dose in the

target region was calculated in cumulative equivalent minutes at 43°C (CEM43) using the

Sapareto and Dewey time-temperature equation [103].

4.2.3 Drug concentration in unheated tissue and VX2 tumours

Tumour-bearing rabbits were administered either lyso-thermosensitive liposomal doxorubicin

(LTLD, Celsion Corporation, Lawrenceville, NJ ) or free doxorubicin (Doxorubicin HCl, Teva

Novopharm, Toronto, ON) at a dose of 2.5 mg/kg body weight diluted with equal parts 5%

dextrose sterile solution. Infusion at 1.2 ml/min into the ear vein during hyperthermia was

initiated once the mean temperature in the target region reached 42°C.

Prior to drug infusion, and at 30 minute intervals afterwards, blood (1.5 ml) was collected

from a catheterized ear artery and transferred to a 2 ml tube containing 167 μl of 0.109 M sodium

citrate to prevent clotting. Plasma was isolated by centrifugation at 4°C for 10 min at 2000 g, and

stored at -20°C.

Approximately two hours after drug infusion, liposomal and free doxorubicin in the

systemic circulation were eliminated by cardiac perfusion with saline under deep anesthesia,

effectively isolating extravasated drug deposited in tissue. For evaluation of tissue drug

concentrations, samples of tumour and adjacent muscle were harvested from the heated and

unheated legs, frozen in liquid nitrogen and stored at -80°C. Samples were also acquired from

the skin, heart, lung, liver, kidney and spleen. Tissue doxorubicin concentrations were measured

by the fluorescence intensity of doxorubicin extracted from homogenized tissue samples as

described in Chapters 2 and 3 [58, 337]. In this study, homogenization of 75 mg tissue samples

was performed using a bead mill homogenizer (Mini-BeadBeater 16, Biospec Products Inc.,

Bartlesville, OK), with 500 μl each of 1 mm and 2 mm zirconia beads.

Doxorubicin concentrations in tissue are summarized by the mean and SD across all

animals. The two-sided Wilcoxon signed-rank test was used to compare drug concentrations in

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heated and unheated tissues for samples of tumour and adjacent muscle. A two-sided Wilcoxon

rank-sum test was used to compare enhancement ratios for thermosensitive liposomes vs. free

doxorubicin. Differences were considered statistically significant for values of p < 0.05.

4.2.4 Drug distribution in the tumour microenvironment

The microregional distribution of doxorubicin with respect to the tumour microvasculature was

measured using quantitative fluorescence microscopy [92]. To improve detection of doxorubicin

fluorescence, two animals were administered increased doses of 8.3 mg/kg of thermosensitive

liposomal doxorubicin during MRI-controlled focused ultrasound hyperthermia. Following post-

treatment imaging, these rabbits were sacrificed without saline perfusion. Heated and unheated

tumours were excised, embedded in optimum cutting temperature compound (OCT, Sakura

Finetek USA, Inc., Torrance, CA), frozen in liquid nitrogen, and stored at -80°C prior to cryostat

sectioning (6 μm).

Blood vessels in tissue sections were recognized by the expression of CD31 membrane

protein on endothelial cells, and doxorubicin distribution was identified by its fluorescence.

Thawed sections were imaged and tiled at 10X and 40X using an Olympus BX50 upright

microscope equipped with a 100 W mercury light source, a Photometrics CoolSnap HQ2 CCD

camera, and a motorized stage. Doxorubicin fluorescence was detected using 3000 ms exposures

with 475 to 495 nm excitation and 589 to 625 nm emission filters prior to immunohistochemical

staining. Following doxorubicin imaging, sections were fixed in acetone for 20 minutes, washed

in PBS, and blocked with a 1:10 dilution of donkey serum for 30 minutes to prevent nonspecific

antibody binding. Sections were then incubated with mouse anti-human CD31 monoclonal

antibodies that show cross-reactivity with rabbit CD31 antigen (ab9498, Abcam Inc., Cambridge,

MA) at a dilution of 1:100 for 1 hour in a humidified chamber, washed in Tris-buffered saline,

and subsequently stained with donkey anti-mouse IgG secondary antibody conjugated to an

Alexa Fluor 488 fluorophore (Jackson ImmunoResearch Laboratories Inc., West Grove, PA) at a

dilution of 1:200 for 1 hour. Finally, slides were cover-slipped with a mounting medium

containing DAPI nuclear dye (Vector Laboratories, Inc., Burlingame CA). Using the same

microscope stage positions used for doxorubicin imaging, anti-CD31 staining of endothelial cells

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was detected using 1000 ms exposures with 475 to 495 nm excitation and 510 to 540 nm

emission filters, and DAPI-stained nuclear DNA localization was detected using 100 ms

exposures (381-392 nm excitation, 420-460 nm emission).

Drug delivery in the tumour microenvironment was evaluated based on the distribution of

doxorubicin in relation to the distribution of tumour blood vessels. Tiled 10X images of CD31-

stained tumour vasculature were thresholded, aligned with the DAPI images based on overlays of

the CD31 and DAPI images, and overlaid on the corresponding images of doxorubicin

fluorescence, which were corrected by subtracting the background reading measured in blank

slide regions. Each tissue section was divided into a grid of 0.5 x 0.5 mm2 regions (0.4

μm2/pixel), excluding areas with staining artifacts and muscle surrounding the tumour. In each

gridded region, the mean doxorubicin fluorescence intensity was recorded, as well as the

microvessel density (measured by the number of vessels per mm2). Doxorubicin fluorescence in

gridded sections of similar vascular density was compared between heated and unheated tumours

using the two-sided Wilcoxon rank-sum test.

4.3 Results

4.3.1 Animals and VX2 tumours

Of 24 rabbits inoculated with bilateral VX2 tumours, 6 were excluded due to tumours that were

poorly positioned or too small to identify on MR imaging. Tumours of at least 8 mm in their

largest dimension were detectable as well-defined regions of heterogeneous signal increase on

T2-weighted MR images. Treated tumours were situated either in the fascia between the biceps

femoris and semimembranosus muscles or infiltrating into one of the two; the largest dimension

of treated tumours ranged between 11 and 28 mm (Table 4.1).

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Figure 4.1: MRI-controlled focused ultrasound hyperthermia in rabbit VX2 tumour. A) Axial T2-weighted MR image depicting experiment setup with tumour-bearing rabbit lying on its side above a degassed water bath containing a mechanically-positioned focused ultrasound transducer. Coronal image planes in which MR thermometry was used to measure temperature during heating are indicated; the middle plane, set at the depth of the ultrasound focus, was used to control treatment. B) Time-averaged MRI temperature measurements in coronal image plane through tumour, demonstrating spatial temperature distribution achieved using MRI-controlled focused ultrasound hyperthermia. Desired temperature was 43°C in the 10 mm diameter targeted region (blue overlay). The focal spot width and scan trajectory are shown. C) Experiment timeline for thermosensitive liposome administration and temperature control. Mean temperatures measured using MR thermometry within the 10 mm diameter target region are shown across three image planes (dark circles), as well as the temperatures that 90% (white circles) and 10% (squares) of the region exceeds. Thermosensitive liposome infusion started when target temperatures reached 42°C (shaded).

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4.3.2 MRI-controlled focused ultrasound hyperthermia in VX2 tumours

The precision of MR temperature measurements used to control heating was 0.25 ± 0.04°C,

measured as the standard deviation of the temperature noise in MR thermometry images. A

linear background magnetic field drift of -0.26 ± 0.06°C/min was corrected online by subtracting

MR temperature measurements made in a mineral oil phantom from temperatures in the rest of

the image. Initial rabbit body temperatures ranged from 34.0°C to 36.7°C and were maintained

within 1°C during treatment.

MRI-controlled focused ultrasound achieved spatially uniform tumour heating, with

temperature elevations localized to the target region (Figure 4.1B) and maintained over the

heating duration (Figure 4.1C). MR temperature measurements in 12 tumour-bearing rabbits are

summarized in Table 4.1. The mean temperature of 42.8°C achieved in the target region closely

matched the goal of 43°C, with 10th and 90th percentile temperatures of 41.4°C and 44.2°C. Drug

infusion began after target region temperatures reached 42°C, which was achieved after a mean

of 6.5 ± 1 minutes. Mean temperatures high enough to achieve drug release from thermosensitive

liposomes (41°C) were localized to an 11.2 ± 1.6 mm diameter region, and maintained above

41°C for 20.3 ± 4.5 minutes in the 10 mm diameter targeted region. The median thermal dose in

the target region was 12.3 CEM43. In 5 rabbits, varying degrees of thermal damage were

observed on post-treatment imaging. Figure 4.2 presents T2-weighted and contrast-enhanced T1-

weighted images for the rabbit that received T10 temperatures of 45.8°C and an accumulated

thermal dose of 26.5 CEM43. In four of the reported experiments, temperature control and

ultrasound exposure were interrupted after less than 20 minutes at the plateau temperature due to

operator error or sudden rabbit motion. Results for six additional rabbits are not reported because

temperature control was interrupted prior to drug infusion.

4.3.3 Drug deposition in VX2 tumours

Plasma doxorubicin concentrations in rabbits treated with hyperthermia during infusion of either

thermosensitive liposomes or free drug are shown in Figure 4.3A. In two early experiments,

blood samples were not collected, and in some animals, blood samples at the latest time point

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Table 4.1: Temperatures measured noninvasively with MR thermometry during MRI-controlled focused ultrasound hyperthermia in rabbit VX2 tumours. Temperatures reported as the temporal mean from when the desired target temperature reached 41°C until the end of heating. Summary data are reported as mean ± standard deviation across all rabbits.

Temperature in target region (Goal: 43°C in 10 mm diameter region for 20 min)

Rabbit mass Tumour

size Tmean T90 T10 Diameter at

41°C Time at 41°C

Thermal Dose

(kg) (mm)a (°C) (°C) (°C) (mm) (MM:SS) (CEM43)

Rabbits administered free doxorubicin (2.5 mg/kg) 2.94 21 x 15 42.6 41.6 43.7 10.3 24:55 16.7 3.21 15 x 10 42.8 41.8 43.7 12.2 24:30 10.7 3.11 14 x 10 42.9 41.2 44.5 9.6 16:30 6.2

3.1 ± 0.1 17 x 12 42.8 ± 0.1 41.5 ± 0.3 44.0 ± 0.5 10.7 ± 1.3 21:58 ± 4:44 10.7 ± 5.3

Rabbits administered lyso-thermosensitive liposomal doxorubicin (2.5 mg/kg) 3.36 28 x 14 42.5 41.6 43.4 10.4 18:00 12.3 3.28 22 x 9 43.2 41.8 44.6 12.5 23:50 8.9 2.95 25 x 7 42.9 41.6 44.3 10.2 13:15 5.2 2.67 18 x 16 42.6 40.6 44.8 11.2 13:35 2.3 2.44 20 x 15 42.7 41.1 44.3 9.2 17:05 12.8 3.06 19 x 11 42.9 40.9 45.8 13.8 25:15 26.5 2.95 18 x 10 42.9 41.8 44.0 14.0 24:35 12.3

3.0 ± 0.3 21 x 12 42.8 ± 0.2 41.3 ± 0.5 44.5 ± 0.8 11.6 ± 1.9 19:14 ± 4:59 12.3 ± 7.7

Rabbits administered lyso-thermosensitive liposomal doxorubicin (8.3 mg/kg) 2.96 13 x 10 42.5 40.7 44.1 9.7 23:35 16.2 3.24 11 x 4 42.6 41.5 43.5 10.9 19:15 13.9

3.1 ± 0.2 12 x 7 42.5 ± 0.1 41.1 ± 0.6 43.8 ± 0.4 10.3 ± 0.9 21:25 ± 3:04 15.1 ± 1.6

3.0 ± 0.3 19 x 11 42.8 ± 0.2 41.4 ± 0.4 44.2 ± 0.7 11.2 ± 1.6 20:22 ± 4:35 12.3 ± 6.3aLargest diameter and perpendicular dimension on axial and coronal T2-weighted images.T90 = temperature exceeded by 90% of points in the target region, T10 = temperature exceeded by 10% of points in the target region.

could not be collected due to clotting in the arterial catheter. For LTLD, an estimated 47% of the

total injected doxorubicin remained in the bloodstream at 30 minutes after the start of infusion

(assuming a total plasma volume of 3.6 ml per 100 g body mass [342]), falling to 15% after 2 h

with a terminal half-life of 88.1 ± 22.8 min (SEM). For free doxorubicin, which has an initial

half-life of less than five minutes [343], rapid plasma clearance left very little drug in the

bloodstream by the time the first sample was collected.

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Figure 4.2: MRI detection of tissue damage in tumour-bearing rabbit treated with hyperthermia and thermosensitive liposomal doxorubicin. The rabbit that received thermal dose of 26.5 equivalent minutes at 43°C is shown. A) T2-weighted images in coronal (top) and axial (bottom) planes through the tumour, acquired prior to treatment (left) and following treatment (right). B) T1-weighted images in the same imaging planes. Gadolinium contrast agent injected immediately before acquiring post-treatment T1-weighted images.

Figure 4.3B summarizes overall drug concentrations in tissue homogenates from rabbits

administered LTLD or free doxorubicin, sampled two hours after the start of drug infusion. The

two drug formulations had similar biodistributions, except for an increased doxorubicin

concentration for LTLD in the liver, where liposomes are known to accumulate [136]. Rabbits

treated with thermosensitive liposomes showed a moderate 4.7-fold enhancement in doxorubicin

deposition in unheated tumours with respect to the surrounding unheated muscle (3.4 ± 1.8 vs.

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Figure 4.3: Plasma and tissue doxorubicin concentrations in tumour-bearing rabbits. A) Blood samples collected prior to and at 30 minute intervals after intravenous administration of lyso-thermosensitive liposomal doxorubicin (LTLD, n = 5) or free doxorubicin (n = 3). Peak concentrations, likely to have occurred at the end of the 6-7 minute infusion (dark shading), could not be measured as this was during MRI-controlled focused ultrasound heating (light shading). Monoexponential fit is shown with 95% confidence interval. B) Biodistribution of doxorubicin two hours after intravenous injection of LTLD (n = 7) or free doxorubicin (n = 3). For both formulations, rabbits were administered 2.5 mg doxorubicin/kg over 6-7 minutes during MRI-controlled focused ultrasound hyperthermia localized to one tumour and its surrounding muscle. Mean ± standard deviation of doxorubicin concentration shown.

0.7 ± 0.2 ng/mg, p = 0.016). In rabbits treated with free doxorubicin, no significant difference

was observed in unheated tumours over unheated muscle (4.9 ± 3.5 vs. 1.1 ± 0.2 ng/mg, p =

0.500), nor for heated over unheated tumours (7.9 ± 1.9 vs. 4.9 ± 3.5 ng/mg, p = 0.250) or

muscle (2.2 ± 1.6 vs. 1.1 ± 0.2 ng/mg, p = 0.250). Following treatment with hyperthermia and

thermosensitive liposomes, doxorubicin concentrations were 26.7 ± 16.2 times higher in heated

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tumours than in unheated tumours (76.3 ± 27.9 vs. 3.4 ± 1.8 ng/mg, p = 0.016), and the thermal

enhancement ratio for thermosensitive liposomal doxorubicin was significantly higher than for

free drug (26.7 ± 16.2 vs. 2.6 ± 2.3 times, p = 0.017). For LTLD, doxorubicin concentrations in

heated muscle surrounding targeted tumours were 22.2 ± 12.3 times higher than in unheated

muscle (15.1 ± 8.8 vs. 0.7 ± 0.2, p = 0.016), but heated tumours had 4.9 ± 1.2 times higher drug

concentrations than heated muscle (76.3 ± 27.9 vs. 15.1 ± 8.8 ng/mg, p = 0.016).

4.3.4 Increased delivery of bioavailable drug in tumour cells

Figure 4.4 provides representative composite images of the microregional distribution of

doxorubicin fluorescence (red) in relation to DAPI-stained cell nuclei (blue) and CD31-stained

blood vessel endothelial cells (green) in heated and unheated VX2 tumours. Both tumours are

from the same animal, harvested 2 hours after intravenous infusion of 8.3 mg/kg doxorubicin in

thermosensitive liposomes. In contrast to the rabbits used for biodistribution measurements in

homogenized samples, no saline perfusion was performed prior to sacrifice. The overall

doxorubicin fluorescence intensity in the heated tumour was higher than in the unheated tumour,

despite having somewhat lower vessel density. Within individual tumours, blood vessel density

was heterogeneous; in heated tumours, doxorubicin fluorescence was higher in regions with

increased vessel density, while in unheated tumours the doxorubicin signal was uniformly low

and indistinguishable from background autofluorescence. Figure 4.5 includes characteristic 0.5 x

0.5 mm2 sections from both highly vascularized (Figure 4.5A, microvessel density approximately

200 vessels/mm2) and poorly vascularized (Figure 4.5B, vessel density 50 vessels/mm2) regions

of the heated and unheated tumours. The doxorubicin fluorescence in the heated tumour

demonstrates a punctate pattern of increased drug accumulation co-localized with DAPI staining

of DNA in the cell, displayed at 40X magnification in Figure 4.6. In heated tumours, nucleus-

specific accumulation of fluorescence from released doxorubicin was consistent even for cells

situated many cell layers from the nearest tumour vessel; in unheated tumours, low levels of

doxorubicin fluorescence were indistinguishable from the background cellular autofluorescence.

This suggests that thermally-triggered release from thermosensitive liposomes increased the

accumulation of bioavailable doxorubicin in the nuclei of tumour cells.

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Figure 4.4: Microregional distribution of doxorubicin in heated and unheated tumours. Rabbits with VX2 tumours implanted in both thighs were treated with MRI-controlled focused ultrasound hyperthermia in one tumour during intravenous infusion of thermosensitive liposomal doxorubicin. Tumours were harvested two hours after infusion and then sectioned, stained and tiled at 10X magnification. Composite images of heated and unheated tumours from one animal display DAPI staining of cell nuclei in blue and background-subtracted doxorubicin fluorescence in red with CD31-stained endothelial cells identifying tumour vessels in green.

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The effect of thermally-triggered release on drug deposition in the microenvironment of

heated and unheated tumours from two rabbits treated with thermosensitive liposomal

doxorubicin is quantified in Figure 4.7A. In 0.5 x 0.5 mm2 regions with matched microvessel

densities of 0 to 400 vessels/mm2, the average doxorubicin fluorescence in sections of heated

tumours was 3.8 to 9.2 times greater than in unheated sections (p < 0.05, two-sided Wilcoxon

rank-sum test). This consistent increase in doxorubicin fluorescence provided by heat-triggered

Figure 4.5: Heterogeneity of microregional distribution of doxorubicin in regions withvarying tumour vascularity. Representative 0.5 mm x 0.5 mm regions of composite imagesfrom heated and unheated tumours showing doxorubicin fluorescence (red), DAPI staining(blue), and CD31 staining of vessel endothelial cells (green) from areas where vessel densitywas relatively high (A) and low (B). The number of vessels per mm2 was measured bycounting distinct regions of CD31-stained endothelial cells.

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Figure 4.6: Accumulation of doxorubicin in the cells of heated and unheated tumours following administration of thermosensitive liposomal doxorubicin. In 40X images of the heated tumour, doxorubicin fluorescence (red) demonstrates co-localization with DAPI staining of cell nuclei (blue) outside of the CD31-stained blood vessels (green). In the unheated tumour, doxorubicin fluorescence was limited to perivascular regions; low levels of doxorubicin fluorescence in the tumour interstitium could not be distinguished from the background.

release, even in regions with low vascular density, suggests improved drug distribution within

the tumour microenvironment. In Figure 4.7B, the uniformity of doxorubicin distribution in the

tumour microenvironment is quantified by the mean ± standard error of doxorubicin fluorescence

intensity in relation to distance to the nearest vessel, averaged across regions containing a

majority of tumour cells in sections from both rabbits. The fluorescence intensity of doxorubicin

in heated tumours decays with distance the way it does with free doxorubicin in mouse tumours

[92], but remains higher than in unheated tumours.

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Figure 4.7: Spatial distribution of doxorubicin fluorescence with respect to vessel density and vessel location in heated and unheated tumours of rabbits administered thermosensitive liposomal doxorubicin. A) Background-subtracted fluorescence intensity of doxorubicin in 500 μm x 500 μm regions of 6 μm frozen sections from two rabbits, classified by the vessel density in those regions. Mean ± SEM is representative of intra-tumour variability based on the aggregate of 675 regions across tumours from two rabbits. * indicates p < 0.05, two-sided Wilcoxon rank-sum test. B) Background-subtracted doxorubicin fluorescence intensity (mean ± SEM of aggregated data from two rabbit tumours) averaged by distance to the nearest CD31-stained pixel across all imaged sections of heated and unheated VX2 tumours from 2 rabbits administered thermosensitive liposomal doxorubicin.

4.4 Discussion In this study, thermally-mediated drug release resulted in enhanced local drug deposition with

specific accumulation of bioavailable drug in the nuclei of tumour cells. These results were

achieved using a commercial formulation of thermosensitive liposomal doxorubicin and

noninvasive MRI-controlled focused ultrasound hyperthermia, in a large animal tumour model.

The large size and heterogeneous vasculature of rabbit VX2 tumours [127, 344] permits testing

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of the robustness of heating techniques, and provides insights on drug delivery in solid tumours

that have distinct regions of high and low vessel density and perfusion [127], similar to tumours

encountered in the clinic.

Our results demonstrate that hyperthermia-mediated drug delivery increases drug

deposition in targeted tissue for a given systemic dose. Administration of thermosensitive

liposomal doxorubicin resulted in a 4.7-fold increase in drug deposition in unheated tumours vs.

surrounding muscle. Localized hyperthermia increased the doxorubicin concentration in heated

tumours and surrounding muscle over unheated tumours and muscle by 26.7 and 22.2 times,

respectively. A 4.9-fold increase was observed in heated tumours over heated muscle. Significant

enhancements were not observed in heated tumours and normal tissue when rabbits were

administered free doxorubicin. With thermosensitive liposomes, enhancements in tumours over

muscle suggest tumour-specific liposome extravasation, while the enhancement in heated over

unheated muscle stresses the importance of localized hyperthermia to prevent unwanted drug

release in normal tissues.

Measurements of plasma doxorubicin concentrations demonstrate an increased

circulation time for doxorubicin in thermosensitive liposomes over free drug [343], suggesting

that prolonged hyperthermia should increase local drug deposition as intact liposomes

continually circulate into heated regions and release their payload.

Similar increases in drug concentration were observed in heated tumours and normal

muscle, despite previous findings demonstrating that the hyperthermia-mediated extravasation of

100 nm liposomes exploited in tumours [130] does not occur in healthy muscle [129].

Furthermore, liposomes were administered during heating, without allowing sufficient time for

passive liposome extravasation and accumulation in targeted tumours. These observations

support the hypothesis proposed by other recent studies [341, 345] that the primary mechanism

for hyperthermia-enhanced drug deposition using thermosensitive liposomes is the triggered

release of bioavailable drug in the vasculature, followed by the subsequent extravasation of

active drug into the tumour interstitium.

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The 26.7-fold enhancement in drug delivery observed with thermosensitive liposomal

doxorubicin in heated tumours is somewhat higher than the 5 to 15-fold enhancements reported

previously [68, 76, 79, 177-180, 184]; however, most of those studies used open-loop, regional

heating techniques in small rodent tumours, with a larger heated volume to body weight ratio.

The recent study of Ranjan et al [341] used a clinical MRI-guided focused ultrasound device and

the same liposome formulation, but reported only a 3.4-fold increase in doxorubicin

concentration between heated and unheated rabbit VX2 tumours. One reason for this discrepancy

is that they heated 4 mm diameter tumour regions to 40-41°C, while we used multi-point

temperature control to heat 10 mm diameter regions to 43°C, with hyperthermic temperatures

extending beyond the tumour along the ultrasound beam. By achieving sufficient temperatures

for triggered drug release (>41°C) in smaller regions, they observed reduced enhancements in

whole tumour homogenates but more specific tumour targeting with respect to adjacent unheated

muscle. In our study, heating to 43°C resulted in thermal damage in some rabbits. This is

expected in rabbit muscle for thermal dose between 5 and 30 CEM43 [106], but underprediction

of thermal damage using thermal dose measurements could have occurred in several ways.

Tumor baseline temperature may have been underestimated if the water bath temperature

exceeded the core body temperature, temperature elevations could have been underestimated due

to averaging across the 5 mm thick image slices, and damage could have been caused by heating

outside of the thermometry image planes. If the thermal dose measurements are correct, these

exposures of up to 30 CEM43 are not likely to cause damage in human tissue [105], and may

provide important clinical benefit, as heat-induced doxorubicin chemosensitization occurs at a

threshold of 43°C [116]. Achieving large increases in tumour drug concentration may be

clinically important, as previous in vitro studies suggest intracellular doxorubicin concentrations

of approximately 60 ng/mg are required for greater than 99% cell kill [54].

Our assessment of the effect of triggered release on the microregional distribution of

doxorubicin by fluorescence microscopy permitted confirmation that enhanced drug

concentrations in heated tumours corresponds to intracellular uptake of bioavailable drug.

Doxorubicin fluorescence co-localized with DAPI staining of cell nuclei throughout the heated

tumour, demonstrating generally increased doxorubicin deposition in comparison with the

unheated tumour. Discrepancies in the degree of doxorubicin fluorescence enhancement in

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heated over unheated tumours between homogenized tissues and microscopy are likely related to

the effects of slide preparation on doxorubicin fluorescence and differences in the excitation

wavelength used for microscopy, as well as the inability to completely isolate doxorubicin

fluorescence from background in unheated tumours. Not performing saline perfusion prior to

tumour harvest for microscopy may also have contributed to an under-estimated enhancement

ratio, although doxorubicin fluorescence may be lower in intact liposomes due to quenching

[326]. These results extend previous qualitative observations of overall doxorubicin distribution

in VX2 tumours following heat-triggered release [341] and quantitative analysis in rabbit thigh

muscle [345].

In agreement with the results of Ranjan et al [341], our measurements of overall

doxorubicin accumulation in homogenized tissue samples from untargeted organs showed a

similar biodistribution in rabbits receiving either free or liposomal drug, with the exception of

increased concentrations in the liver for LTLD, where liposomes are expected to accumulate.

Additional experiments using fluorescence microscopy in untargeted organs would be useful in

determining whether drug accumulation in those tissues resulted from the extravasation of

circulating liposomes or from intravascular drug release.

MR thermometry data demonstrated that focused ultrasound, controlled by quantitative

MR images of tissue temperature, can be used to achieve temporally and spatially uniform mild

hyperthermia in heterogeneously-perfused tumours. The mechanically-steered system used in

this study is a cost-effective preclinical approximation of electronically-steered clinical MRI-

guided focused ultrasound systems [255], providing flexibility in customizable scanning

trajectories and multi-point temperature control algorithms in a research setting. However, to

reduce the effects of transducer motion on MR thermometry, we used a single-element

transducer with a relatively high f-number (weakly focused) that could heat the tumour while

being located far from the imaging plane [337]. The use of an f-number 2 transducer resulted in

heating being spread out along the axis of the ultrasound beam, causing drug release in muscle

along the beam path. Our control algorithm could easily be applied with existing clinical devices,

in which phased array transducers would allow better control of the ultrasound field, while multi-

plane MR thermometry could improve safety and temperature control [300, 346].

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4.5 Conclusion Our study contributes to a growing body of literature that demonstrates localized, image-guided

drug deposition in implanted tumours, using MRI-controlled focused ultrasound hyperthermia to

achieve noninvasive thermally-mediated drug release from thermosensitive liposomes [179, 180,

341, 345, 346]. Heat-triggered release of doxorubicin from liposomes in the tumour vasculature

allowed bioavailable drug to accumulate in the tumour and localize at its site of lethal activity in

the nuclei of tumour cells. Our study complements recent work by achieving large increases in

doxorubicin concentration in tumours for a given systemic dose, and providing quantitative

analysis of drug distribution in the tumour microenvironment. Clinically, enhancing the

accumulation and penetration of anticancer drugs in solid tumours could increase cell kill in each

chemotherapy treatment cycle and limit the opportunity for tumours to repopulate and gain drug

resistance between cycles, thus providing patients with locally advanced solid tumours a better

chance of relapse-free survival.

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Chapter 5 Conclusions and Future Work

5.1 Summary of findings The treatment of solid tumours using anticancer drugs is impeded by barriers to drug transport in

the tumour microenvironment that limit the delivery of cytotoxic intracellular concentrations of

chemotherapeutic agents in all tumour cells without causing toxicity in normal tissues. This

thesis explores the use of thermally-mediated drug delivery to localize chemotherapy to solid

tumours, and to uniformly increase drug concentrations in tumour cells. It describes the

preclinical development of MRI-controlled focused ultrasound hyperthermia for noninvasive,

precisely controlled, localized heating of centimeter-scale targets deep within the body, and its

use in achieving doxorubicin release from thermosensitive liposomes in vivo.

Chapter 2 evaluated the feasibility of using MRI-controlled focused ultrasound and

thermosensitive liposomes to achieve thermally-mediated localized drug delivery in vivo. Results

were reported from ten rabbits, where an ultrasound focus was scanned in a circular trajectory to

heat 10-15 mm diameter regions in normal thigh to 43°C for 20-30 minutes. MRI thermometry

was used for closed-loop feedback control to achieve temporally and spatially uniform heating.

Closed-loop control of FUS heating using MRI thermometry achieved temperature distributions

with mean, T90 and T10 of 42.9°C, 41.0°C and 44.8°C, respectively, over a period of 20

minutes. Lyso-thermosensitive liposomal doxorubicin was infused intravenously during

hyperthermia. The fluorescence intensities of homogenized samples from heated and unheated

thigh regions were used to calculate the concentration of doxorubicin in tissue. Doxorubicin

concentrations were significantly higher in tissues sampled from heated compared to unheated

regions of normal thigh muscle (8.3 versus 0.5 ng/mg, mean per-animal difference = 7.8 ng/mg,

P < 0.05, Wilcoxon matched pairs signed rank test). The results showed the potential of MRI-

controlled focused ultrasound hyperthermia for enhanced local drug delivery with temperature-

sensitive drug carriers.

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Chapter 3 evaluated the feasibility of achieving image-guided drug delivery in bone by

combining MRI-controlled focused ultrasound hyperthermia and systemic administration of

temperature-sensitive liposomes. Hyperthermia (43°C, 20 minutes) was generated in 10 mm

diameter regions at a muscle-bone interface in nine rabbit thighs using focused ultrasound under

closed-loop temperature control using MR thermometry. Thermosensitive liposomal doxorubicin

was administered systemically during heating. Heating uniformity and drug delivery were

evaluated for control strategies with the temperature control image centered 10 mm (four rabbits)

or 0 mm (five rabbits) from the bone. Simulations estimated temperature elevations in bone.

With the ultrasound focus and MR temperature control plane 0 mm and 10 mm from the bone

interface, average target region temperatures were 43.1°C and 43.3°C; numerically-estimated

bone temperatures were 46.8°C and 78.1°C. With the 10 mm offset, temperatures in the target

region were accurately controlled, but strong ultrasound absorption at the bone surface resulted

in thermal ablation outside of the focus; numerically-estimated muscle temperatures were 66.1°C

at the bone interface. Drug delivery was quantified using the fluorescence of doxorubicin

extracted from bone marrow and muscle, and compared between treated/untreated thighs using

the one-sided Wilcoxon signed-rank test. Localized increases in doxorubicin concentration

occurred in image-targeted regions of heated vs. unheated marrow (8.2-fold, p = 0.002) and

muscle (16.8-fold, p = 0.002). Enhancement occurred for 0 mm and 10 mm offsets, suggesting

localized drug delivery in bone is possible with both hyperthermia and thermal ablation. The

results of this work demonstrate that MRI-controlled focused ultrasound can achieve localized

hyperthermia in bone, for noninvasive image-guided drug delivery in bone with temperature-

sensitive drug carriers.

Chapter 4 addressed the inability of anticancer drugs to reach all cancer cells in solid

tumours at sufficient concentration to cause cytotoxicity. Hyperthermia-triggered release of

drugs from thermosensitive liposomes can increase tumour drug concentration, but tumour-

specific drug delivery requires precise temperature control, and effects on microregional

distribution of anticancer drugs in tumours are unknown. Here we evaluated thermally-triggered

release of doxorubicin in a rabbit tumour model by comparing free vs. thermosensitive liposomal

doxorubicin administered systemically during magnetic resonance imaging (MRI)-controlled

focused ultrasound hyperthermia. Twelve rabbits with a transplanted VX2 tumour in each thigh

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had a 10 mm diameter region in one tumour heated to 43°C using focused ultrasound with

temperature control by MRI thermometry. Delivery of doxorubicin to tumours and normal

tissues was quantified by fluorescence in tissue homogenates, and by fluorescence microscopy.

Using thermosensitive liposomal doxorubicin (2.5 mg/kg), doxorubicin concentrations in heated

tumours were 26.7 times higher than in unheated tumours (n = 7, p = 0.017, two-sided Wilcoxon

signed-rank test). There was no significant enhancement with free doxorubicin in heated vs.

unheated tumours (n = 3, p = 0.50). With thermosensitive liposomes (8.3 mg/kg), fluorescence

microscopy demonstrated increased doxorubicin fluorescence in heated vs. unheated tumours,

co-localized with nuclear staining throughout the tumour. Localized image-guided delivery of

high concentrations of doxorubicin to cancer cells was achieved noninvasively in implanted

tumours with temperature-sensitive drug carriers and a preclinical MRI-controlled focused

ultrasound hyperthermia system.

Overall, this work described how focused ultrasound hyperthermia could be used to

achieve localized delivery of chemotherapeutic agents from temperature-sensitive drug carriers.

With temperature sensitive liposomes, heat-triggered release achieved drug concentrations in

heated tumours of 76 ng/mg, more than 20 times that of unheated tumours, and 10 times that of

tumours in rabbits given the same dose of free drug. Fluorescence microscopy of heated and

unheated tumours from rabbits given thermosensitive liposomal doxorubicin demonstrated that at

the same time point, the high overall drug concentrations in heated tumours corresponded to

uniformly elevated microregional drug distributions, co-localized with DAPI-stained cell nuclei.

This pattern of increased intracellular doxorubicin fluorescence remained high even several cell

layers away from the nearest CD31-stained tumour blood vessel, suggesting strong uptake even

in poorly-perfused tumour regions.

The importance of achieving and maintaining high tumour drug concentrations has been

demonstrated by in vitro studies with the anticancer drug doxorubicin in human lung cancer cell

lines [54]. The rate of cellular uptake is increased with high extracellular drug concentrations,

and intracellular doxorubicin concentrations are closely related to tumour cell death, with

approximately 60 ng/mg causing 99% cell death [54]. With localized and sustained hyperthermia

combined with thermosensitive liposomes, intact liposomes continually flow into the tumour,

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where they rapidly release drug in the tumour vasculature. Released doxorubicin can then

extravasate into the tumour, acting as a localized high-dose infusion of the active drug. This

results in a strong concentration gradient that drives drug diffusion through the tumour

microenvironment, and prolongs exposure to high extracellular drug concentrations for increased

intracellular uptake and cell kill. Additionally, maintaining a high rate of cellular uptake could

overwhelm drug export from the cytoplasm in multi-drug resistant cells. To further increase

exposure time for greater intracellular uptake, thermally-induced vascular shutdown after mild

hyperthermia could be used to prevent drug washout as systemic concentrations decline.

These results support the use of focused ultrasound hyperthermia with temperature

sensitive liposomes in the treatment of solid tumours, and can be used to guide hyperthermia

timing and drug dosing in translation of this technology to the clinic. Our MRI-controlled

hyperthermia system is a prototype device designed for use in preclinical studies, but models

focused ultrasound systems that are approved for use in humans. Studies in the same animal

model are underway to further investigate the effects of localized release on tumour regression,

and human trials have been approved for combining thermally-mediated drug delivery with

focused ultrasound thermal ablation in the palliative treatment of bone tumours.

5.2 Limitations

5.2.1 Quantification of doxorubicin fluorescence

In this thesis, two techniques were used to measure doxorubicin in tissue. Calibrated

fluorometric measurements of doxorubicin fluorescence were used to quantify the total amount

of doxorubicin present in blood plasma and homogenized tissue samples. Fluorescence

microscopy was used to evaluate the pattern of doxorubicin uptake with respect to tumour cells

and tumour vasculature, providing qualitative information on the microregional distribution, and

relative measurements of overall fluorescence intensity as a function of vessel density and

distance to the nearest vessel. It is important to recognize several limitations of these

measurement techniques.

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5.2.1.1 Doxorubicin fluorometry in homogenized tissue samples

Fluorometric measurements of overall doxorubicin concentrations in tissue samples were useful

in determining the degree to which localized release increased drug accumulation in targeted and

untargeted regions. However, several assumptions are implied and sources of error arise at each

step of this technique including tissue harvest, sample preparation, drug extraction, and

fluorescence detection.

Tissues were harvested approximately two hours after drug infusion, at which point most

unbound doxorubicin will have washed out of interstitial tissues, and most liposomes in the

systemic vasculature will have cleared from circulation. To more completely isolate extracellular

plus intracellular doxorubicin in tissue, rabbits were sacrificed by cardiac perfusion with saline,

thereby clearing any remaining drug-carrying liposomes from the bloodstream. Having only one

time point limits our ability to characterize tumour drug permeation and washout; however,

harvesting tissue after drug washout and saline perfusion provides a reasonable measure of the

amount of doxorubicin that will remain bound in the tissue to have an antitumour effect. Another

issue related to tissue harvest is that only small tissue samples were collected and the weight of

entire organs was not recorded upon dissection. Collection of entire organs or at least recording

the weight would have provided more detailed information on biodistribution, such as the

percent injected dose delivered to each tissue type.

To expedite sample processing, stored samples from several animals were homogenized

on the same day. To protect doxorubicin in harvested tissue from the photobleaching effects of

light and metabolic degradation, samples were frozen in liquid nitrogen and stored in the dark at

-80°C prior to homogenization and doxorubicin extraction. The homogenization process breaks

down tissue structure and ruptures cells, allowing doxorubicin extraction from cellular organelles

and liposomes with organic solvents. In our early studies, tissue homogenization was performed

in acidified ethanol extraction buffer using glass tissue grinders. This time-intensive method

required approximately 1 hour of manual labour per sample. Later, use of an automated bead

mill homogenizer permitted homogenization in the same extraction buffer with less time and

effort, without affecting measured fluorescence intensity. While the bead mill homogenized 16

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samples in only three minutes, careful weighing and preparation of samples for homogenization

kept the total time per sample at about 10 minutes.

The extraction buffer used in our studies was first used for anthracyclines by Bachur,

who concluded on the use of 0.3N hydrochloric acid in 50% ethanol after comparing several

organic solvents for their ability to extract bound drug and their effect on the drug's emitted

fluorescence spectrum [58]. Subsequent studies have had success using the same [60, 68, 70] and

other [80, 347] organic solvents for doxorubicin extraction. A two-step process using chloroform

and silver nitrate can be used to separate free drug from doxorubicin bound tightly to DNA and

RNA in the cell nucleus [73, 76, 348]. However, other investigators found similar overall

extraction efficiencies for both techniques [65]. Fluorometry of doxorubicin in isolated tumour

cell nuclei may provide more accurate quantification of bioavailable versus sequestered and

liposomal drug [83].

Fluorescence detection introduces another set of limitations. The fluorometer used in this

study uses a halogen lamp with excitation and emission filters specific to peaks of the

doxorubicin excitation and emission spectra. The repeatability of fluorescence intensity

measurements is sensitive to temperature, light source stability, and the accuracy of the

doxorubicin fluorescence calibration in various tissues. To limit these sources of error,

fluorometry was always performed under the same ambient light conditions, with doxorubicin

calibration standards prepared each day, normalized by doxorubicin calibration curves in tissue

homogenates with known doxorubicin concentrations. Reported measurements were the mean of

recordings made on three separate days. On each day, recorded fluorescence intensities were the

mean of three readings. To account for temporal variation in source intensity, temperature, or

photobleaching, all 12-24 samples in a batch were read once before making the second reading

for all samples.

Fluorescence intensity measurements made in these fixed wavelengths include signal

from doxorubicin metabolites, which exhibit similar emission spectra to doxorubicin and are

present in varying proportions in different organs [55]. Fluorescence from drug metabolites can

be isolated by measuring full fluorescence spectra in tissue samples using high-performance

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liquid chromatography with fluorescence detection, and doxorubicin content can also be

measured by mass spectrometry [349].

The overriding limitation of drug quantification in bulk tissue samples is that it cannot

determine the microregional distribution of doxorubicin in tissue, and thus has limited ability to

predict whether increased overall drug concentrations will translate into increased cell kill within

the complex tumour microenvironment.

5.2.1.2 Fluorescence microscopy in frozen tumour sections

To gain insight on the microregional distribution achieved using triggered drug release, we used

a previously described wide-field fluorescence microscopy technique [92], quantifying the

fluorescence intensity of doxorubicin with respect to CD31-stained tumour vessels and DAPI-

stained cellular structures.

While fluorescence intensity gradients are related to changes in drug concentration,

several factors can influence fluorescence intensity in tiled microscope images, making it

difficult to make absolute measurements of drug concentration in tumour cells. Intensity

variations are introduced by variations in light source intensity, the ambient environment, and

section thickness, as well as fluorescence artifacts on the slide and folds or rips in tumour

sections. These sources of error were addressed by careful preparation and selection of

microscope slides for analysis, as well as acquisition and comparison of background

fluorescence images before and after each tiled acquisition. Photobleaching is another important

consideration when imaging doxorubicin fluorescence; to minimize this source of variation in

fluorescence intensity, short exposure times were used to adjust the focus, calibrate the stage,

and define the tiled region for each acquisition. Inside the cell, doxorubicin accumulates in acidic

organelles where decreased pH may alter the fluorescence emission; however, previous work in

Dr. Tannock’s group suggests this may not have a significant effect on doxorubicin fluorescence

[95]. At the two hour time point, thermally-mediated delivery resulted in a pattern of primarily

intracellular doxorubicin fluorescence; it is possible that due to quenching, these fluorescence

intensities underestimate high concentrations of DNA-bound drug [350].

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In addition to these variations in fluorescence intensity, our analysis of doxorubicin

fluorescence intensity with respect to distance to CD31-stained tumour vessels in 8-10 μm thick

tumour sections introduces important geometric limitations when characterizing drug delivery in

the three-dimensional tumour microenvironment. The inability to detect blood vessels above and

below the imaged section overestimates the distance of doxorubicin in each pixel to the vessel it

diffused from. Conversely, many CD31-stained tumour vessels are actually not perfused, and

doxorubicin fluorescence in nearby pixels likely results from diffusion from further vessels.

Additionally, after stitching of tiled images, spatial offsets became apparent between images

acquired with the three different filter sets. Images were registered by manually overlaying

stitched images to align fluorescence artifacts and tumour edges, but in some cases a spatially

varying offset remained. However, by averaging over large tumour regions and performing

careful background subtractions, these geometric effects should not strongly influence the overall

distribution profiles.

Neither the microscopy nor the fluorometric doxorubicin quantification techniques used

in this thesis can determine whether elevated drug concentrations in unheated tumours is the

result of systemic free doxorubicin concentrations, or the accumulation of intact liposomes in the

tumour. One approach to differentiating free and liposomal drug would be to repeat the

experiment using fluorescein-labeled liposomes, and imaging their distribution directly. This has

recently been performed in a mouse window chamber tumour model, demonstrating that during

hyperthermia, intact fluorescein-labeled liposomes remain inside the tumour vessels, while heat-

triggered release allows extensive diffusion of doxorubicin out of the vessel, through the tumour

interstitium and into tumour cells [351].

5.3 Further studies: The effect of localized drug delivery on antitumour efficacy

5.3.1 Study motivation and design

The results of this thesis demonstrate that localized drug release using MRI-controlled focused

ultrasound hyperthermia achieves 10-20 fold increased DOX deposition in VX2 tumours, with

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high intracellular DOX uptake throughout the tumour. Based on these locally enhanced

intracellular drug concentrations, a study has been initiated to investigate the effect of thermally

mediated drug delivery on tumour response in the same tumour model. Specifically, this study

will address the question of whether localized release of doxorubicin from thermosensitive

liposomes using MRI-controlled focused ultrasound hyperthermia improves antitumour efficacy

in the rabbit VX2 tumour model.

Rabbits bearing a VX2 tumour in one thigh will be administered lyso-thermosensitive

liposomal doxorubicin with or without 20 minutes of hyperthermia at 42.5°C using MRI-

controlled focused ultrasound. This study will use the Philips clinical MR-HIFU system instead

of a preclinical device, using hyperthermia strategies that can be applied directly to humans.

Treatment efficacy will be evaluated based on tumour progression, with tumour volume

measured weekly using contrast-enhanced T1-weighted MRI. Indicators of early doxorubicin-

related toxicity will be myelosuppression and decreases in body weight. Dermal toxicity related

to liposome and drug accumulation in the skin, including alopecia and skin lesions, will be

monitored clinically. The major long-term toxicity of doxorubicin is degeneration of cardiac

muscle, which will be monitored by elevated serum levels of creatine kinase and lactate

dehydrogenase, and later identified using histology.

5.3.2 Doxorubicin antitumour efficacy and toxicity

The doxorubicin dose for this study was identified after reviewing the literature for preclinical

efficacy and toxicity of doxorubicin, liposomal doxorubicin, and temperature-sensitive liposomal

doxorubicin.

With single intravenous doses of free doxorubicin in rabbits with VX2 thigh tumours

[352], a dose of 1 mg/kg had no effect on tumour growth compared to controls treated with

saline, while 2 mg/kg achieved a tumour growth plateau at 4 weeks. A dose of 3 mg/kg was still

well-tolerated and achieved measurable tumour regression in 10 out of 12 rabbits, with complete

response in 6 rabbits. At 4 mg/kg, rabbits experienced severe toxicity including alopecia, limb

wasting, muscle atrophy, and weight loss.

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In a toxicity study, rabbits received a doxorubicin dose of 1.2 mg/kg per week,

administered every 3-5 days for 3-11 weeks, with 6 months of follow-up [353]. Bone marrow

suppression was identified in the blood counts after 3 weeks of treatment, recovering after 2

weeks off-treatment. Rabbits with cumulative doses of greater than 10 mg/kg, sacrificed after 3

to 6 months of follow-up, showed progressive degeneration of cardiac muscle. In these animals,

earlier blood tests had sustained elevation of creatine kinase (a common indicator of muscle and

myocardial damage) and lactate dehydrogenase (which is elevated during tissue breakdown, but

may also be elevated due to tumour growth). Congestive heart failure was observed in 2 of 6

rabbits in the maximum cumulative dose group (24 mg/kg).

Using non-temperature sensitive liposomal doxorubicin, which circulates for several days

and accumulates slowly in the tumour, incidence of doxorubicin-related progressive

cardiotoxicity was reduced even after cumulative doses of 21 mg/kg (1 mg/kg every 5 days) [51].

With free doxorubicin, severe cardiomyopathy was observed in 10 of 15 rabbits at 14 mg/kg,

with congestive heart failure causing the death of 5 animals. With liposomal drug, minor

cardiomyopathy was detected on histology for only 3 of 15 animals at 14 mg/kg, and 1 of 10 at

21 mg/kg. However, liposomal doxorubicin accumulation in the skin caused significant dermal

toxicity. After 5 weeks (cumulative dose of 7 mg/kg), open sores in the paws, as well as redness

and painful crusting of the skin under the limbs forced treatment to be postponed for 14 days,

which was enough time for most animals to recover; however, skin lesions led to early sacrifice

in 6 of 25 rabbits. Reversible, but dose-limiting dermal toxicity has also been documented in

humans for the same liposome formulation [154].

The pharmacokinetics and biodistribution of thermosensitive liposomal doxorubicin are

somewhere between free and liposomal doxorubicin. Increased drug delivery using triggered

release should improve antitumour efficacy over free doxorubicin for a given systemic dose,

while not circulating long enough for dangerous accumulation in the skin. In our study, we plan

to use a dose of 1.67 mg/kg. This is similar to the 2 mg/kg minimum effective single dose of free

doxorubicin in rabbit VX2 thigh tumours, shown to achieve stable disease but not regression.

Converting this dose to other species using normalized body surface area [324], our rabbit dose

can be compared to thermosensitive liposomal doxorubicin doses used in previous mouse and rat

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studies of tumour regression. Our rabbit dose of 1.67 mg/kg is equivalent to a standard mouse

dose of 6.67 mg/kg, slightly above the 5 mg/kg that demonstrated strong tumour regression in

human squamous cell carcinoma xenografts [76], JC murine mammary carcinoma [68], and

EMT-6 tumours [180], and quite similar to the 6-7 mg/kg that showed varying degrees of

improved survival over thermosensitive liposomal doxorubicin or heat alone in several other

mouse tumour lines [177]. Our dose is also equivalent to a standard rat dose of 3.34 mg/kg,

lower than the 5 mg/kg that showed improved survival in a rat fibrosarcoma model [79]. In that

study, doxorubicin without heat, heat alone, and thermosensitive liposomal doxorubicin alone

were all similar to the saline control. Our single doxorubicin dose provides a much lower

cumulative drug or liposome exposure than has been shown to cause dermal or cardiac toxicity.

Our rabbit dose is also equivalent to a standard human dose 20 mg/m2, the lowest dose used in

Phase I/II breast cancer trials that employed less selective external microwave hyperthermia

techniques [185]. With focused ultrasound, we hope to use improved hyperthermia localization

and control to achieve greater antitumour effect with a low injected dose.

5.3.3 Using a clinical HIFU system for MRI-controlled hyperthermia

Preclinical studies of tumour response with thermosensitive liposomal doxorubicin have only

been undertaken in small animal models. Most have used water bath heating techniques [76, 79,

177, 354], with one study using poorly monitored heating by pulsed ultrasound exposures [68].

Our study will provide efficacy data for hyperthermia-mediated drug delivery with

thermosensitive liposomes in a non-rodent tumour model. VX2 rabbit tumours are more similar

to human cancers than mouse or rat tumours in terms of size, intervening tissue interfaces, and

their heterogeneous blood flow, which make it difficult to achieve noninvasive, uniform heating

localized to prescribed target regions. The rabbit VX2 tumour model thus allows us to use an

accurately localized, clinically relevant hyperthermia technique to investigate the efficacy of

thermally-mediated drug delivery. Using the clinical Philips MR-HIFU system, our study will

have similar temperature elevations and localization of heating as can be expected in humans.

The Philips MR-HIFU system uses a 256-element phased array therapeutic ultrasound

transducer built into an MRI patient bed [255]. It electronically steers the 1.2 or 1.4 MHz focus

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along the boundary of circular trajectories with 4, 8, 12 and 16 mm diameter (Figure 5.1). For

mild hyperthermia, we use a modified version of the binary feedback control algorithm

developed to rapidly create a uniform region of irreversible thermal damage for ablation of

uterine fibroids [280]. Our algorithm has a goal of maintaining a mean temperature of 42.5°C for

20 minutes, similar to one recently presented by Partanen et al [346]. During the initial heat-up

phase, the interior circles are heated to 39°C. Next, the outer circle is heated until it reaches the

lower range of the target temperature band (in this case 42.3°C). In the maintenance phase, zero

power is applied until the mean temperature of the pixels along the boundary of any circle drops

below 42.3°C, at which point the underheated circle is sonicated. When sonicating, power is shut

off again if the mean temperature in any circle exceeds the upper limit of the target temperature

band (42.5°C), or if the maximum temperature exceeds a safety threshold of 44.0°C.

Figure 5.1: Sonication and MR thermometry setup using Philips MR-HIFU system. Left:Electronically-focused phased array transducer steers focus along concentric circular pathsto achieve uniform heating of a targeted region. Right: Orthogonal view depicting 4, 8, 12and 16 mm subtrajectories and relative size of ultrasound focus. MR thermometry slicesare overlaid, with three images across the focal plane, two orthogonal images along thedirection of ultrasound propagation, and one additional slice positioned at the skin.

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5.4.4 Early results

Figure 5.2 displays contrast-enhanced T1-weighted planning images and temperature maps taken

from the treatment planning software during MRI-controlled hyperthermia in a rabbit VX2

tumour. The clinical software reconstructs a contiguous set of images (in this case axial) into

three orthogonal planes, for target definition as well as transducer translation and rotation to

avoid heating untargeted tissues like the bone. During sonication, MR temperature maps are

acquired in six planes centered on the target every 10 seconds: three images across the beam

through the target, two orthogonal images along the beam path, plus one image in a user-defined

plane to detect skin heating where the beam enters the thigh (Figure 5.1). In this example, a mean

temperature of 41.5°C was achieved over 20 minutes in the 12 mm diameter target region

centered on the tumour, although overheating to temperatures of 44°C occurred in muscle

beyond the tumour causing thermal damage. This demonstrates some of the technical challenges

of using a clinical system for preclinical studies. The clinical transducer is made to heat large

regions deep within the human body, operating at low frequency (1.2-1.4 MHz) with a long focal

depth (12 cm). With a scanned focus, the heated region is about twice as large along the beam as

it is across, heating a 24 mm long region for a 12 mm diameter cell. This typically spans the

entire rabbit thigh, making it difficult to localize heating along the beam. In Figure 5.2,

transverse temperature maps show high temperature readings not only in the tumour, but also at

the skin where the beam enters and exits the thigh, and in the muscle between the tumour and the

inner thigh.

Figure 5.3 shows follow-up contrast-enhanced T1-weighted images of the first two

rabbits in our study, demonstrating a marked difference in tumour progression between the rabbit

treated with thermosensitive liposomal doxorubicin alone (top) compared to the rabbit treated

with thermosensitive liposomal doxorubicin and MRI-controlled hyperthermia (bottom). This

promising initial data point will be combined with data from a planned 5 more rabbits in each

treatment group to overcome innate variability in VX2 tumour growth.

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Figure 5.2: Treatment planning and MRI-controlled hyperthermia in a rabbit VX2 tumour using clinical MR-HIFU system. Top: 3D reformatting of contrast-enhanced T1-weighted planning images. Ultrasound beam overlay used to guide transducer positioning. Bottom: MR temperature maps acquired after 15 minutes of controlled heating.

Figure 5.3: Rabbit VX2 tumour response following thermally-mediated drug delivery using thermosensitive liposomal doxorubicin (TLD) and MRI-controlled focused ultrasound. Contrast-enhanced T1-weighted images acquired at treatment planning, and at 1, 2, and 3 weeks post-treatment to monitor tumour growth. Top: Rabbit treated with TLD alone. Bottom: Rabbit treated with TLD plus MRI-controlled heating.

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5.4 Future directions and clinical applications

5.4.1 Thermally-mediated drug delivery using other therapeutic agents

and delivery vehicles

The most thoroughly studied thermosensitive drug carriers to date have been lyso-

thermosensitive liposomes containing the widely used chemotherapeutic agent doxorubicin. This

formulation has been instructive in preclinical development and gaining clinical approval, but

several other therapeutic payloads, drug carriers, and drug release mechanisms may prove to

have greater efficacy against specific cancers and in other diseases.

Doxorubicin is successfully used as a single agent or as part of combination

chemotherapy protocols in the treatment of numerous cancer types [355]. Its fluorescence

provides a convenient means for evaluating drug delivery in preclinical studies, and its increased

cytotoxicity in combination with hyperthermia makes it a natural choice for use in thermally-

mediated drug delivery. However, other anticancer agents such as cisplatin, bleomycin,

carmustine, melphalan, and cyclophosphamide offer greater cytotoxicity for specific cancer

types, while also exhibiting additive or synergistic effects with heat [114]. While each payload

drug has unique vascular permeability, diffusivity, and rate of cellular uptake, these differences

can be accounted for in drug carrier design, creating the possibility of using thermally-mediated

drug delivery as a generalized approach to localized chemotherapy for solid tumours.

As cancer genomes for the various tumour types are revealed, new genetically targeted

agents are being developed to act on the pathways responsible for each of the hallmarks of

cancer, classifying cancers not by organ, but by the specific mutations identified in their genome.

However, drugs that target genetic mutations still need to reach tumour cells in concentrations

sufficient to cause cytotoxic or antiproliferative effects, and may benefit from combination with

thermally-mediated drug delivery as described in this thesis. Localized release of high

concentrations of cytotoxic drugs could be administered in combination with molecular targeted

agents, providing indiscriminate cell kill to augment the targeted agents' cytostatic effect on

drug-resistant cells, for robust treatment of heterogeneous tumour populations. Alternatively,

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hyperthermia-triggered release could be used to achieve a localized infusion of small molecule

targeted agents capable of efficient extravasation and diffusion, thereby increasing their

therapeutic index. For targeted antibodies, triggered release could be used to deliver

pharmacological modifiers of the tumour microenvironment, such as enzymes that degrade the

fibrous extracellular matrix, pH-modifiers that sensitize tumour cells to chemotherapy, p-

glycoprotein inhibitors designed to reverse multi-drug resistance, or to deliver anti-vascular

agents with the intent of normalizing tumour vasculature to sensitize tumours to chemotherapy

and radiation. By delivering cytotoxic, molecular-targeted, or tumour-sensitizing agents,

thermally mediated drug delivery could facilitate and generalize anticancer drug therapy by

overcoming variations in pharmacokinetics between patients and organs.

In addition to the numerous formulations of temperature-sensitive liposomes [175, 180,

301, 356], promising thermosensitive drug carriers include polymer micelles [357], lipid

nanoparticles [358], and soluble copolymers [165]. Polymeric micelles are self-assembling

bubbles of 10-100 nm diameter, formed from amphiphilic block copolymers whose hydrophobic

regions aggregate together to form a hydrophobic core enclosed by a hydrophilic corona [359].

Micelles are effective carriers of water-insoluble agents; in some cases, their small size enables

intracellular delivery, and thermosensitive formulations offer rapid drug release [357]. Similarly,

solid lipid nanoparticles can be used to achieve extracellular or intracellular release of water-

soluble drugs [167]. Yet another approach is to bind drugs to copolymer molecules that are

soluble in the bloodstream but become insoluble and aggregate upon heating, causing localized

drug accumulation at the heated site [165]. Some of these nanocarriers are designed to release

their payload in the vasculature, others in the extracellular space, and still others are designed to

enter the cell before drug release. While these drug delivery systems offer their own drug release

profiles, each approach would benefit from the strategies for localized, controlled, noninvasive

focused ultrasound hyperthermia as described in this thesis.

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5.4.2 Imageable nanoparticle drug carriers

By incorporating imaging contrast agents into temperature-sensitive nanoparticle drug carriers,

thermally-triggered release can be used to make a one-time measurement of absolute temperature

during thermal therapy [360]. This absolute temperature information is especially useful in

tissues where uncertain baseline temperatures, high lipid content, or magnetic field disturbances

from motion or respiration preclude use of the proton resonance frequency shift thermometry

technique. For example, when gadolinium is loaded into temperature-sensitive liposomes, intact

liposomes provide minimal contrast enhancement due to limited access of gadolinium to free

water. When tissue temperatures cross the melting temperature of the drug carrier’s bilayer

lipids, gadolinium is released from the liposome, and its interaction with bulk water causes

enhanced longitudinal relaxation that can be detected on T1-weighted MR images [246, 361,

362]. This concept can be expanded to include two temperature-sensitive mechanisms in a single

imageable nanoparticle, allowing detection of low and high temperature boundaries to guide

mild heating without thermal coagulation [358]. In this approach, gadolinium is loaded into a 10-

14 nm hydrogel nanoparticle that exposes the contrast agent to free water only below the

hydrogel melting temperature, providing an “on-off” contrast enhancement transition at an upper

temperature threshold. These hydrogel nanoparticles are then incorporated into a 100 nm solid

lipid nanoparticle, which releases the hydrogel and exposes its contrast agent to free water at the

melting temperature of the lipid nanoparticle, creating an “off-on” transition at a selected lower

temperature threshold [358].

In addition to monitoring temperature, temperature-sensitive macromolecular constructs

loaded with both chemotherapeutic agents and MR contrast agent have been used to achieve

thermally-mediated image-guided drug delivery [163, 179, 354, 363-365]. Upon heating, both

the drug and the tracer are released; if they have similar extravasation, diffusion, and binding

properties, this results in modified image contrast in the vicinity of the released drug. For

example, lyso-thermosensitive liposomes formulated to release both doxorubicin and gadolinium

when heated to 41°C have demonstrated localized T1-weighted signal increase in heated regions

of rat and rabbit tumours after MR-guided FUS hyperthermia [179, 365]. Imageable drug carrier

development has been supported by the use of MR imaging techniques capable of both

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temperature measurement and detection of gadolinium contrast enhancement [366]. This

combination, though still in preclinical development, would enable the clinician to noninvasively

assess and direct drug accumulation and release in an approach known as “drug dose painting”

[79]. The control strategies developed in this thesis could be applied to achieve “dose-controlled”

drug delivery by automatically modifying the heating trajectory (regions heated to 41°C) based

on contrast agent release detected by dynamic T1-weighted imaging.

5.4.3 Clinical applications for thermally mediated drug delivery

Localized drug delivery using ultrasound hyperthermia has the potential to improve the

specificity of chemotherapy for any solid tumour that can be heated, and to which liposomes can

enter via the vasculature. This section discusses several specific cancer types to which the

hyperthermia strategies and drug formulation used in this thesis are most readily applied in

humans. With further developments in hyperthermia strategies and thermosensitive drug carriers,

applications will expand to include currently intractable diseases such as pancreatic cancer and

glioblastoma.

A clinical trial has already been approved to add thermosensitive liposomes to the

thermal ablation of painful bone metastases using MR-HIFU [367]. This primarily palliative

application represents a proving ground for the safety of focused ultrasound mediated drug

delivery in the clinic. Results from this trial would form a basis for the translation of this

technology to the more urgent indication of unresectable primary tumours in pediatric

osteosarcoma, rhabdyosarcoma, and Ewing sarcoma [368]. In these young patients, surgery is

disfiguring and sometimes debilitating; radiation and systemic chemotherapy including

doxorubicin, while demonstrating antitumour effect, pose significant risks of secondary

malignancies and long-term toxicity. Non-ionizing focused ultrasound hyperthermia could be

used to heat children’s bone and soft tissue tumours with the control algorithms developed in this

thesis to achieve localized exposure to doxorubicin, which is already used to treat these tumours.

Our lab is developing intracavitary devices for MRI-controlled ultrasound hyperthermia

of locally advanced colorectal cancer [369] and thermal coagulation of prostate cancer [275].

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These diseases are less sensitive to doxorubicin, but thermosensitive nanoparticles loaded with

more appropriate cytotoxic agents are currently under development, such as 5-fluorouracil

thermosensitive polymer nanoparticles [370] or oxaliplatin for rectal cancer, and doxetaxel-

loaded thermosensitive polymer hydrogels [371] or micelles [372] for prostate cancer. Like

doxorubicin, these agents have also been demonstrated to have heat-enhanced effect [373, 374].

Another indication being studied at our institution is the use of hyperthermia in the treatment of

head and neck cancer. Head and neck cancers are responsive to cisplatin [375], the cytotoxicity

of which is strongly enhanced by hyperthermia [376, 377]. It’s possible that localized delivery

from recently developed lyso-thermosensitive liposomal cisplatin could enhance drug effect or

reduce cisplatin-induced nephrotoxicity [378].

Motion-tracking MRI thermometry and focused ultrasound delivery techniques are also

being developed to apply thermally-mediated drug delivery to the treatment of pancreatic cancer,

which presents the dual challenges of respiratory motion and sonication through the ribs [379-

381]. MR thermometry sequences are also being optimized for imaging near fat and in the

presence of lung motion [382-385], to be used with new transducer designs for treating breast

cancer [386, 387]. For patients with locally advanced breast cancer undergoing preoperative

chemotherapy including doxorubicin, thermally-mediated drug delivery could be used to lower

systemic toxicity, thereby enabling more aggressive neoadjuvant chemotherapy to render

inoperable tumours resectable.

The immediate clinical application of thermally-mediated drug delivery has the promise

to increase extracellular drug concentrations and improve intracellular drug uptake for improved

therapeutic index in a variety of cancers, primarily in the context of locally advanced or locally

recurrent solid tumours, and localized early-stage disease. However, the ability to localize

exposure to chemotherapy is also the main limitation of this approach, as chemotherapy is

typically used for the systemic eradication of widespread metastatic disease. While commercially

available thermosensitive liposomes exhibit approximately 10% drug release in the circulation at

normothermic temperatures, low levels of circulating free drug are unlikely to achieve cytotoxic

concentrations in unheated metastases. Furthermore, localized tumours are often adequately

treated by surgery or radiotherapy. This restricts thermally-mediated drug delivery to use in cases

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of localized cancers, where chemotherapy has a cytotoxic effect on tumour cells, but surgery

poses undue risk to the patient or results in disfigurement, and radiation is either ineffective or

poses unacceptable risk for late toxicity and the development of secondary malignancy. In many

such cases, thermal ablation is a valid treatment option; thermally-mediated drug delivery can be

combined with slow (10 to 20 minutes) thermal coagulation of the clinical target volume,

achieving local tumour destruction with enhanced cell kill at the thermal margins to reduce the

risk of local recurrence over ablation alone. In cancers where a tissue sparing approach would be

preferred in order to preserve quality of life, mild hyperthermia can be used to achieve localized

high-dose chemotherapy, for local tumour control with minimal tissue damage and systemic

toxicity. Using MRI-controlled focused ultrasound hyperthermia to trigger drug release from

thermosensitive liposomes, these cancers could be effectively treated in a precise, noninvasive

manner, expanding the range of treatment options for patients with localized or early-stage solid

tumours.

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