On the Permeabilisation and Disruption of Cell Membranes by Ultrasound and Microbubbles

Raffi Karshafian

A thesis submitted in conformity with the requirements for the degree of Doctor of Philosophy Graduate Department of Medical Biophysics University of Toronto

Raffi Karshafian Doctor of Philosophy Thesis Department of Medical Biophysics, University of Toronto, Sunnybrook Health Sciences Centre, S639-2075 Bayview Avenue Toronto, Ontario M4N 3M5 Canada

On the Permeabilisation and Disruption of Cell Membranes by Ultrasound and Microbubbles Raffi Karshafian Doctor of Philosophy Department of Medical Biophysics, University of Toronto, 2010

Therapeutic efficacy of drugs depends on their ability to reach the treatment target. Drugs that exert their effect within cells are constrained by an inability to cross the cell membrane. Methods are being developed to overcome this barrier including biochemical and biophysical strategies. The application of ultrasound with microbubbles increases the permeability of cell membranes allowing molecules, which otherwise would be excluded, to enter the intracellular space of cells; a phenomenon known as sonoporation. This thesis describes studies aimed at improving our understanding of the mechanism underpinning sonoporation and of the exposure parameters affecting sonoporation efficiency. Cancer cells (KHT-C) in suspension were exposed to ultrasound and microbubbles – total of 97 exposure conditions. The effects on cells were assessed through uptake of cell-impermeable molecules (10 kDa to 2 MDa FITC-dextran), cell viability and microscopic observations of the plasma membrane using flow cytometry, colony assay and electron microscopy techniques. Sonoporation was a result of the interaction of ultrasound and microbubbles with the cell membrane. Disruptions (30-100 nm) were generated on the cell membrane allowing cell impermeable molecules to cross the membrane. Molecules up to 2 MDa in size were delivered at high efficiency (~70% permeabilisation). Sonoporation was short lived; cells re-established their barrier function within one minute, which allowed compounds to remain inside the cell. Following uptake, cells remained viable; ~50% of sonoporated cells proliferated. Sonoporation efficiency depended on ultrasound and microbubble exposure conditions. Microbubble disruption was a necessary but insufficient indicator of ultrasound-induced permeabilisation. The exposure conditions can be tailored to achieve a desired effect; cell permeability of ~70% with ~25% cell death versus permeability of ~35% with ~2% cell death. In addition, sonoporation depended on position in the cell cycle. Cells in later stages were more prone to being permeabilised and killed by ultrasound and microbubbles. This study indicated that sonoporation can be controlled through exposure parameters and that molecular size may not be a limiting factor. However, the transient nature may necessitate that the drug be in close vicinity to target cells in sonoporation-mediated therapies. Future work will extend the investigation into in vivo models.

ii To my wife Taleen and children Karnie, Tro and Meghrie

To my parents, who gave up everything to give us a better life

iii Acknowledgements

Dr. Peter N. Burns, my scientific mentor, provided an excellent environment to satisfy my scientific curiosity and the guidance in becoming an independent scientist, for which I am very grateful. I thank my supervisory committee members, Drs. Dick Hill, Kullervo Hynynen and Brian Wilson for their guidance and scientific discussions. Many colleagues including research engineers and graduate students provided support throughout my graduate years. Most notably, Ross Williams, Sanya Samac, Anoja Giles, Dr. Emmanuel Cherin, Dr. Peter D. Bevan and Kasia Harasiewicz. Steven Doyle (Microscopy Imaging Lab, University of Toronto) helped with electron microscopy preparations. I have also benefited from Drs. Gregory J. Czarnota and David E. Goertz scientific knowledge and insightful discussions. Finally, I am grateful for the continuous support from my family and friends throughout my academic path. This work was supported in part by the Canadian Institutes of Health Research (CIHR) Doctoral Award (3 years) and the Ontario Graduate Scholarship in Science and Technology (OGSST).

iv TABLE OF CONTENTS ABSTRACT...... ii Acknowledgements...... iv TABLE OF CONTENTS...... v LIST OF TABLES...... x LIST OF FIGURES...... xi

Chapter One Drug Delivery and Sonoporation 1.1 Introduction...... 1 1.2 The problems of anticancer therapeutic agents...... 2 1.3 Biological barriers to anticancer agents...... 2 1.3.1 Cell membrane barrier...... 3 1.4 Delivery strategies for therapeutic agents...... 4 1.4.1 Intracellular delivery strategies...... 4 1.5 Ultrasound and microbubbles in imaging and therapy...... 5 1.5.1 Physics of ultrasound...... 5 1.5.2 Ultrasound in imaging and therapy...... 6 1.5.3 Ultrasound microbubble contrast agents...... 6 1.5.4 Ultrasound imaging with microbubbles...... 8 1.5.5 Ultrasound therapy with microbubbles...... 9 1.6 Sonoporation...... 9 1.6.1 Sonoporation Efficiency...... 10 1.6.2 Mechanism of Sonoporation...... 12 1.7 Sonoporation applications...... 13 1.8 Thesis outline...... 14

Chapter Two Sonoporation by ultrasound-activated microbubble contrast agents: Effect of acoustic exposure parameters on cell membrane permeability and cell viability 2.0 Abstract...... 16

v 2.1 Introduction...... 16 2.2 Methods...... 17 2.2.1 In vitro cell model...... 17 2.2.2 Ultrasound exposure system...... 18 2.2.3 Ultrasound microbubble agent...... 19 2.2.4 Microbubble size distribution: Coulter Counter...... 20

2.2.5 Reversible permeability and PI-viability: PR and VPI...... 20

2.2.6 Therapeutic Ratio: TRR...... 21 2.2.7 Experiments: Ultrasound exposure parameters...... 22 2.3 Results...... 24 2.3.1 Ultrasound exposure parameters...... 24 2.3.2 Optimisation of sonoporation: Therapeutic Ratio...... 30 2.3.3 Relationship between permeability and viability...... 34 2.3.4 Relationship between permeability and microbubble disruption...... 35 2.4 Discussion...... 36 2.4.1 Optimisation of sonoporation...... 36 2.4.2 Mechanism of sonoporation...... 38 2.4.3 Role of microbubble disruption...... 39 2.4.4 Limitations of the study...... 39 2.5 Conclusions...... 41

Chapter Three Microbubble mediated sonoporation of cells: Clonogenic viability and influence of molecular size on uptake 3.0 Abstract...... 42 3.1 Introduction...... 43 3.2 Methods...... 43 3.2.1 In vitro cell model...... 43 3.2.2 Ultrasound exposure system...... 43 3.2.3 Ultrasound microbubble agent...... 44

3.2.4 Reversible permeability and PI-viability: PR and VPI...... 44

3.2.5 Clonogenic Assay: CRP and CUP...... 44 3.2.6 Experiments...... 45

3.2.7 Therapeutic Ratio: TRR and TRC...... 46

vi 3.3 Results...... 47 3.3.1 Microbubble agent...... 47 3.3.2 Uptake of different molecular size markers...... 49 3.3.3 Viability of reversibly permeabilised cells...... 50

3.3.4 Therapeutic Ratio: TRR and TRC...... 53 3.4 Discussion...... 54 3.4.1 Reversible and viable permeabilisation...... 54 3.4.2 Microbubble agent...... 54 3.4.3 Limitations of the study...... 55 3.5 Conclusions...... 56

Chapter Four Mechanism of cell sonoporation: Generation of transient sub-micron disruptions on the plasma membrane by ultrasound and microbubbles 4.0 Abstract...... 57 4.1 Introduction...... 58 4.2 Methods...... 58 4.2.1 In vitro cell model...... 58 4.2.2 Ultrasound exposure system...... 58 4.2.3 Ultrasound microbubble agent...... 58

4.2.4 Reversible permeabilisation and PI-viability: PR and VPI...... 58 4.2.5 Electron microscopy...... 59 4.3 Results...... 59 4.3.1 Intracellular uptake of FITC-dextran...... 59 4.3.2 Disruption of plasma membrane...... 61 4.4 Discussion...... 65 4.4.1 Mechanism of sonoporation...... 65 4.4.2 Disruption induced by ultrasound-activated microbubbles...... 66 4.4.3 Limitations of the study...... 67 4.5 Conclusions...... 67

vii Chapter Five Dependence of sonoporation on cell cycle: Enhanced effect during later stages - Work in progress 5.0 Abstract...... 68 5.1 Introduction...... 69 5.2 Methods...... 69 5.2.1 In vitro cell model...... 69 5.2.2 Ultrasound exposure system...... 69 5.2.3 Ultrasound microbubble agent...... 69 5.2.4 Reversible permeability and PI-viability: Cell cycle phase...... 69

5.2.5 Measures of cell sensitivity: DPI and PE...... 70

5.2.6 Therapeutic Ratio: TRP...... 70 5.3 Results...... 71 5.3.1 Cell PI-viability...... 71 5.3.2 Reversible permeability...... 71

5.3.3 Therapeutic Ratio: TRP...... 74 5.3.4 Intracellular uptake...... 75 5.4 Discussion...... 76 5.4.1 Relationship between sonoporation and cell cycle...... 76 5.4.2 Limitations of the study...... 76 5.5 Conclusions...... 77

Chapter Six Summary and Future Directions 6.1 Summary and Conclusions...... 78 6.2 Future Directions...... 81 6.2.1 Sonoporation of in vitro endothelial cells...... 82 6.2.2 Sonoporation of in vivo CAM blood vessels...... 84 6.3 Final Comments...... 88

References...... 89

viii APPENDIX...... 99 A. Awards...... 99 B. Publications...... 99

ix LIST OF TABLES

Table 2.1: A summary of optimum and maximum therapeutic ultrasound exposure parameters for the conditions considered in this set of studies...... 33

x LIST OF FIGURES

Figure 1.1: Ultrasound exposure parameters: Pulse centre frequency, peak negative pressure, pulse duration, pulse repetition frequency and insonation time...... 11 Figure 1.2: (a) Sonoporation phenomenon: The application of ultrasound increases the permeability of cell membranes and allows delivery of molecules which othewise would be excluded. (b) Sonoporation studies described in this thesis, which aim at improving our understanding of sonoporation, include the investigation of: (1) the effect of ultrasound exposure conditions, (2) the effect of microbubble exposure conditions, (3) the effect of molecular size on sonoporation-mediated uptake, (4) cell viability, the ability of cells to proliferate, following uptake, (5) the mechanism underpinning sonoporation and (6) the effect of cell cycle phase on sonoporation...... 15 Figure 2.1: A schematic diagram of the ultrasound exposure apparatus. Cells are placed within the chamber and exposed to different acoustic conditions...... 18 Figure 2.2:The ultrasound field of the 500 kHz unfocused transducer with 32 mm element diameter focused at 85 mm...... 19 Figure 2.3: The effects of peak negative pressure and pulse frequency on cell permeability and viability. The percentages of permeabilised and viable cells are shown with respect to pressure at centre frequencies of 500kHz (solid), 2MHz (dash) and 5MHz (dash-dot), where P is the number of permeabilised cells which are still viable, and V is the number of viable cells. Cell permeability increases and viability decreases with acoustic pressure, and cell permeability decreases and viability increases with centre frequency. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-3.5MPa, f=0.5-5MHz, T=120s, and Definity (3.3% volume concentration)...... 25 Figure 2.4: The effects of peak negative pressure and pulse frequency on Definity microbubbles ranging in size between 1 μm and 8 μm. (a) The total number of microbubbles disrupted (per ml) exposed to different peak negative pressures and pulse centre frequencies. The size distribution of Definity (bubbles/ ml) exposed to (b) 500 kHz, (c) 2 MHz and (d) 5 MHz for various acoustic pressures. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-2.5MPa, f=0.5-5MHz, T=120s, and Definity (3.3% volume concentration)...... 26 Figure 2.5: The effects of pulse duration and peak negative pressure on cell permeability and viability. The percentages of permeabilised and viable

cells are shown with respect to pulse duration, where PR is the number of

cells are shown with respect to pulse repetition frequency, where PR is the number

of permeabilised cells which are still viable, and VPI is the number of viable cells: 125 kPa (solid); 246 kPa (dash); 570 kPa (dash-dot). Cell permeability

xi increases and viability decreases with pulse repetition frequency and acoustic pressure. Exposure conditions: PD=32μs, PRF=10-3000Hz, Pneg=0-570kPa, f=500kHz, T=120s, and Definity (3.3% volume concentration)...... 28 Figure 2.7: The effects of pulse duration (dash) and pulse repetition frequency 2 (solid) at constant acoustic energy (ESPPA=3.1 J/cm ) on cell permeability and viability. The percentages of permeabilised and viable cells are shown with

respect to duty cycle, where PR is the number of permeabilised cells which are

still viable, and VPI is the number of viable cells. Cell permeability increases and viability decreases with pulse duration and pulse repetition frequency at constant acoustic energy. Exposure conditions: PD=0-32μs, PRF=10- 3000Hz, Pneg=570kPa, f=500kHz, T=3-900s, and Definity (3.3% volume concentration)...... 29 Figure 2.8: The effects of insonation time and peak negative pressure on cell permeability and viability. The percentages of permeabilised and viable

cells are shown with respect to insonation time, where PR is the number of

permeabilised cells which are still viable, and VPI is the number of viable cells: 125 kPa (solid); 246 kPa (dash); 570 kPa (dash-dot). Cell permeability increases and viability decreases with insonation time and acoustic pressure until they reach a plateau. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz, T=0-120s, and Definity (3.3% volume concentration)...... 31 Figure 2.9: The therapeutic ratio, the proportion of cells in which reversible permeability increase is induced compared to those that are killed by the same exposure, is shown for various (a) peak negative pressures and pulse frequencies, (b) pulse duration and peak negative pressure, (c) pulse repetition frequency and peak negative pressure, (d) pulse duration and pulse repetition frequency at constant acoustic energy density, and (e) insonation time and peak negative pressure...... 32 Figure 2.10: The therapeutic ratio, the proportion of cells in which reversible permeability increase is induced compared to those that are killed by the same exposure, is shown as a function of acoustic energy density (J/cm2) for all

exposure conditions. TRR values range from ~0.2 to 8.8 at acoustic energy 2 densities of 1-10 J/cm . A similar TRR of ~3.5 can be achieved at energy densities ranging from 0.5 to 3000 J/cm 2 ...... 33 Figure 2.11: The data for permeabilised cells and viable cells for all exposure conditions are shown in (a). Cell membrane permeability increases to 70%

while viability decreases to 75%. (b) A contour graph of TRR as a function of

permeability and viability indicates areas of high and low TR R values...... 34 Figure 2.12: The number of permeabilised and viable cells with respect to the fractional number of remaining bubbles is shown here for all exposure

conditions, where PR is the number of permeabilised cells which are still

viable, and VPI is the number of PI-viable cells. Bubble disruption appears to be a necessary but insufficient condition to permeabilise cell membranes...... 35 Figure 3.1: Sonoporation of KHT-C cells with 10 kDa FITC-dextran. (a) Flow cytometry fluorescence intensity of FITC-dextran versus propidium iodide (PI). The upper-left hand quadrant shows cells which are permeabilised and viable. Positive PI staining indicates cell death. At these conditions, 78±2%

xii of the cells were permeabilised, with only 18±1.5% killed. (b) Fluorescent confocal microscopy of a permeabilised cell of 1mm thickness showed the 10 kDa FITC-dextran molecules inside the cell. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=120s, and Definity (3.3% v/v)...... 47

Figure 3.2: Cell permeability (PR) and PI-viability (VPI) with Definity agent;

microbubble concentration was varied from 0 to 13.2% v/v. PR, cells stained with 70 kDa FITC-dextran and unstained with PI (PI-viable) molecules, increased with microbubble concentration and reached a maximum level of 71±2% % at 3.3% v/v, beyond which it decreased. The 3.3% v/v was identified

as the optimal concentration of Definity. PI V decreased with microbubble concentration. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz and. T=120s...... 48

Figure 3.3: Cell permeability (PR) and PI-viability (VPI) with Optison agent;

microbubble concentration was varied from 0 to 13.2% v/v. PR, cells stained with 70 kDa FITC-dextran and unstained with PI (PI-viable) molecules, increased with microbubble concentration and reached a maximum level of 44±1% % at 6.7% v/v, beyond which it decreased. The 6.7% v/v was

identified as the optimal concentration of Optison. VPI decreased with Optison microbubble concentration. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz and T=120s...... 48 Figure 3.4: Uptake of different molecular weight FITC-dextran markers with Definity agent. Comparable uptake of the FITC-dextran marker of 10 kDa, 70 kDa, 500 kDa and 2 MDa in size was achieved at fixed acoustic pressure.

Reversible permeability (PR) increased with acoustic pressure for all marker

sizes. Maximum PR of 78±2% was achieved with 10 kDa FITC-dextran at 570 kPa. Exposure conditions: Definity (3.3% v/v), PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s...... 49 Figure 3.5: Uptake of different molecular weight FITC-dextran markers with Optison agent. Comparable uptake of the FITC-dextran marker of 10 kDa, 70 kDa, 500 kDa and 2 MDa in size was achieved at fixed acoustic pressure.

Reversible permeability (PR) increased with acoustic pressure for all marker

sizes. Maximum PR of 56±4% was achieved with 10 kDa FITC-dextran at 570 kPa. Exposure conditions: Optison (6.7% v/v), PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s...... 50 Figure 3.6: Clonogenicity of sonoporated cells. The number of cell colonies formed by reversibly permeabilised and unpermeabilised cells, normalized with respect to untreated control, is shown against acoustic pressure for two microbubble agents (Definity at 3.3% v/v and Optison at 6.7% v/v).

Clonogenicity of reversibly permeabilised cells with Definity (D-CRP) and

Optison (O-CRP) decreased to 30-38% and 48-51%, respectively. Clonogenicity of unpermeabilised cells following ultrasound and microbubble treatment

remained high; clonogenicity of unpermeabilised cells with Definity (D-CUP)

and Optison (O-CUP) was 70-98% and 89-91%, respectively. Clonogenicity of reversibly permeabilised cells was lower compared to clonogenicity of unpermeabilised cells. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0- 570kPa, f=500kHz and T=120s...... 51

xiii Figure 3.7: Permeabilised and clonogenically viable cells. Viable permeability

(PCV), the number of cells which are permeabilised and clonogenically

viable, and clonogenic viability (VC), the number of clonogenically viability

(VC), are shown with two agents (Definity and Optison) and four acoustic pressures (0-570 kPa). Maximum viable permeability of 22% was achieved

with both agents; D-PCV and O-PCV refer to viable permeability with Definity and Optison, respectively. A higher clonogenic viability was achieved with

Optison compared to Definity; D-VC and O-VC refer to clonogenic viability with Definity and Optison, respectively. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s...... 52 Figure 3.8: Therapeutic ratio of sonoporated cells with viability assessed using

propidium iodide (PI) and clonogenic assay. The therapeutic ratios, TRR and

TRC, the number of permeabilised cells divided by the number that are killed by the same exposure, are shown with two agents (Definity and Optison) and

four acoustic pressures (0-570 kPa). A higher TRR, viability assessed with

PI, was achieved with Definity (D-TRR) than Optison (O-TRR). However, a

higher TRC, viability assessed with colony assay, was achieved with Optison

(O-TRC) than Definity (D-TRC). Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s...... 53 Figure 4.1: Uptake of FITC-dextran during sonoporation and one minute following

termination of the treatment. Cell permeability (PR), the percentage of cells

stained with FITC-dextran and unstained with PI, and cell viability (VPI), the percentage of cells unstained with PI, are shown for four conditions: Untreated, ultrasound alone, ultrasound and microbubbles with FITC-dextran added 60 s before treatment, and ultrasound and microbubbles but with FITC-dextran added 60 s after terminating the ultrasound treatment. Intracellular delivery of FITC-dextran is achieved only when present during treatment with ultrasound

and microbubbles. VPI was reduced to 80% by ultrasound and microbubbles. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v...... 60 Figure 4.2: Pore-like structures and ruffles on scanning electron microscopy (SEM) images of cells. SEM image of untreated cell shows numerous ruffles (a), which are absent on ultrasound and microbubble treated cells (c); acquired at 5 000x and 10 000x magnifications. Pore-like structures are observed on untreated cells (~100 nm) (b) and on treated cells (~300 nm) (d); acquired at 40 000 x magnification. The arrow indicates a structure that resembles a pore. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v...... 61 Figure 4.3: High-resolution transmission electron microscopy (TEM) images at 80 000 times magnification of (a) control cells, (b) cells treated with ultrasound alone and fixed immediately, (c) cells treated with ultrasound and microbubbles and fixed immediately; and (d) cells treated with ultrasound and microbubbles and fixed one minute following ultrasound exposure. Disruptions on the plasma membrane are observed in (c) which are absent in (a), (b) and (d). Black arrow indicates a disruption on the plasma membrane. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity=3.3%

xiv v/v...... 63 Figure 4.4: Histogram of disruption size generated on the membrane of cells by ultrasound and microbubbles. Disruptions in the range of 30-to-100 nm were observed with few of the disruptions as large as 400 nm. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v. Based on analysis of 12 TEM images of cells treated with ultrasound and microbubbles and fixed immediately following termination of the treatment No disruptions were observed on the TEM images of cells before treatment and one minute following termination of the treatment...... 64

Figure 5.1: PI-viability (VPI) in the whole population, and G1, S and G2/M phases.

VPI decreased with cell cycle stage and acoustic pressure...... 72

Figure 5.2: The median and the 25th and 75th percentiles of PI-dead cells (DPI) in different phases of the cell cycle. The number of cells stained with PI molecules were normalized with respect to the number of cells in controls in the whole

population, G1, S and G2/M phases. Statistically significant differences was

measured in the number of cells killed in the later stages (G2/M) compared to

early stages (G1)...... 72

Figure 5.3: Reversible cell permeability (PR) measured at varying acoustic pressures

(0, 100, 125 and 570 kPa) in the whole population, and G1, S and G2/M phases. The number of cells stained with 70 kDa FITC-dextran and unstained with PI

molecules, indicating reversible permeabilisation (PR) are shown in the major phases of the cell cycle...... 73 Figure 5.4: The median and the 25th and 75th percentiles of the number of

permeabilised cells (PE) in different phases of the cell cycle and acoustic

pressures. PE, the ratio of the number of reversibly permeabilised cells with respect to the number of viable cells remaining following treatment in whole

population, G1, S and G2/M phases. A statistically sifnificant higher PE levels were achieved at the later stages...... 73

Figure 5.5: Therapeutic ratios, TRP, the number of permeabilised cells divided by

the number that are killed by the same exposure, of the whole population, G1,

S and G2/M phases are shown. Lower TRP values were achieved in the later stages of the cell cycle...... 74 Figure 5.6: Fluorescence intensity of FITC-dextran delivered to the intracellular

space of cells in the whole population, G1, S and G2/M phases are shown. Higher intensities were achieved at the later stages of the cell cycle and higher acoustic pressures...... 75 Figure 6.1: Conclusions of this thesis: 1) Mechanism of sonoporation is the generation of transient sub-micron disruptions on cell membrane, 2) Sonoporation can be controlled by exposure conditions, 3) Cell-impermeable molecules can be taken up at high efficiency, 4) Microbubble disruption is a necessary but insufficient indicator of ultrasound-induced permeabilisation, and 5) Long- term viability of sonoporated cells is possible...... 80 Figure 6.2: The effect of peak negative pressure on cell viability of two endothelial cell lines. The percentages of viable cells are shown with respect to peak negative pressure: Cell viability decreases with peak negative pressure. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-250kPa, f=0.5MHz,

xv insonation time=2 minutes, and Definity 3.3% v/v...... 83 Figure 6.3: Therapeutic ratio is shown for varying peak negative pressure. The therapeutic ratio is optimum at 246kPa peak negative pressure. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-250kPa, f=0.5MHz, insonation time=2 minutes, and Definity 3.3% v/v...... 83 Figure 6.4: (a) The chorioallantoic membrane (CAM) of a fertilized chicken egg at day 15 with the shell removed. Blood vessels are optically visible. (b) A photograph depicting the ultrasound setup with the microscope unit on top. (c) A schematic diagram of the ultrasound exposure system...... 84 Figure 6.5: (a) Control photograph of CAM vasculature perfused with microbubbles and fluorescent marker prior to ultrasound treatment. (b) Perfused vasculature 30 seconds into treatment; no noticeable treatment effects is observed. (c) CAM vasculature after five minutes of treatment; noticeable effect of vessel constriction is observed. (d) A photograph of CAM after treatment. Vessel constriction and disappearance of capillaries is apparent compared with control photograph...... 86 Figure 6.6: Fluorescence microscopy images of excised CAM layer vasculature. (a) Control vessels from area untreated by ultrasound. (b) Fluorescence image of control vessels. (c) Ultrasound treated area. (d) Fluorescence image of ultrasound treated area...... 87

xvi LIST OF SYMBOLS AND ABBREVIATIONS c Speed of sound in water CAM Chorioallantoic membrane

CRP Clonogenicity of reversibly permeabilised cells

CUP Clonogenicity of unpermeabilised

DPI PI-dead cells

ESPPA Spatial-peak-pulse-average acoustic energy f Pulse centre frequency

ISPPA Spatial-peak-pulse-average intensity

PCV Viable permeability PD Pulse duration

PE Permeabilisation efficiency PI Propidium Iodide Pneg Peak negative pressure

PR Reversible permeability PRF Pulse repetition frequency

Prms Root mean square of peak negative pressure PRP Pulse repetition period T Insonation time

TRC Therapeutic ratio based on PRC and VC

TRP Therapeutic ratio with cell cycle phase

TRR Therapeutic ratio based on PR and VPI

VC Clonogenic viability

VCNTL Mean viability of untreated control

VPI PI-viable cells

VUP Unpermeabilised cells r Density of surrounding medium τ Exposure time

xvii CHAPTER ONE Drug Delivery and Sonoporation

“We can’t solve problems by using the same kind of thinking we used when we created them” Albert Einstein (1879-1955)

1.1 Introduction The therapeutic efficiency of biologically active molecules that exert their effect within the cell is constrained by the inability of complex molecules to cross the cell membrane (Jain 2001, Larkin et al. 2008). The cell membrane restricts delivery of therapeutic molecules such as chemotherapeutic drugs and genetic materials to the intracellular space of cells (Minchinton and Tannock 2006). Biochemical to biophysical strategies are being investigated to overcome this barrier (Alper 2002, Gupta et al. 2005, Wu et al. 2008). These methods have in common the requirement to transport agents to their site of action at sufficient concentrations in a safe and reproducible manner with the aim of improving therapeutic efficacy while simultaneously reducing toxic side effects associated with the treatment. This dissertation investigates a phenomenon, known as sonoporation, where the application of ultrasound with microbubbles temporarily increases the permeability of cell membranes and allows molecules to cross the otherwise-impermeable membrane and enter the cell. Sonoporation-mediated therapeutic applications are being developed for cancer, cardiovascular and blood-brain-barrier limited treatments (Bekeredjian et al. 2005, Ferrara et al. 2007, Hynynen 2008). However, it remains a challenge to achieve consistent and controllable delivery, in part because the process of sonoporation is not well understood.

Chapter One 1 1.2 The problems of anticancer therapeutic agents Cancer affects many people and is a major cause of death in humans (Jemal et al. 2008). It is a complex disease in which cells become abnormal, lose their function and divide without control. Solid cancers manifest themselves through the development of in situ lesions. These lesions remain dormant and cannot grow beyond a few millimeters in size in the absence of angiogenesis – the formation of new blood vessels (Folkman 2007). The tumour vasculature which continuously expands with the growing tumour allows cancer to metastasize by permitting cancer cells to invade other tissues in the body and form secondary lesions (Bergers and Benjamin 2003, Folkman 1995). Treatment of cancer still remains a challenge (Kim and Tannock 2005). The treatment must kill all cells within a cancer that are capable of regenerating it. Clinically, solid tumours are surgically removed when possible. The cancer is generally treated with radiation and/or anticancer therapeutic agents targeting cancerous cells and/or their supporting vasculature (Jain 2005, Neri and Bicknell 2005). The impact of an anticancer therapeutic agent, which can be chemical, biological or genetic, depends on the rate and extent to which their constituents penetrate tissues and cells to reach their intended target (Minchinton and Tannock 2006). Treatment effectiveness is limited by toxic side effects exerted by therapeutic agents on tissues and cells not associated with the disease (Allan and Travis 2005). Ideally, therapeutic agents would be delivered entirely to diseased tissue, sparing healthy tissues. However, anticancer agents generally lack specificity – they are unable to selectively target cancer cells and blood vessels associated with the disease (Orive et al. 2003). Furthermore, treatment effectiveness is limited in solid tumours due to inadequate delivery of therapeutic molecules to the disease site as a result of biological barriers in cancerous tissues (Jain 1999, Jain 2001, Minchinton and Tannock 2006).

1.3 Biological barriers to anticancer agents The biological barriers that molecules of a therapeutic agent encounter before reaching their intended target depend on the mode of administration. Generally, anticancer therapeutic agents are administered intravenously in the systemic circulation. Although administering agents through the oral route can achieve a similar effect, bioavailability is usually a concern (Orive et al. 2003). Therapeutic agents must pass through the stomach and intestines before entering the blood compartment. However, once therapeutic agents enter the blood compartment, they are distributed throughout the body where they can interact with diseased as well as healthy tissues

Chapter One 2 and organs. For example, chemotherapeutic agents damage rapidly dividing normal cells of the bone marrow and the gastrointestinal tract. Large therapeutic molecules generally are cleared through the reticuloendothelial system (RES) – the cellular system responsible for clearance of foreign material from the body. These unwanted interactions can cause dose-limiting toxicity and may prevent the therapeutic agent from reaching the intended target at the required concentration. Molecules of pharmacological compounds reaching tumour tissues encounter additional barriers before reaching the site of their therapeutic action. Therapeutic molecules are heterogeneously distributed within a tumour tissue in part because the transport dynamics within and across the vasculature is highly abnormal and inefficient. The morphology of the tumour vasculature is chaotic with poorly organized architecture. It contains tortuous vessels, shunts, vascular loops and large avascular areas (Tozer et al. 2005). The blood flow is irregular and sluggish. Furthermore, the blood vessels of tumours are leaky and tumours usually lack a functional lymphatic system resulting in increased levels of interstitial fluid pressure (IFP). The elevated IFP inhibits delivery and distribution of the therapeutic molecules (Heldin et al. 2004). Additionally, the abnormal nature of tumour vasculature gives rise to hypoxic and acidotic areas reducing the effectiveness of radiotherapy and many anticancer drug therapies (Brown and Wilson 2004, Jain et al. 2007).

1.3.1 Cell membrane barrier The therapeutic efficacy of biologically active molecules that exert their therapeutic action inside the cell is constrained by the inability of complex molecules to cross the cell membrane (Jain 2001). The cell membrane (plasma membrane) separates cells from their environment and restricts the passage of molecules. It consists of lipids that are arranged in a bilayer structure, forming a hydrophilic exterior and a hydrophobic interior of ~5 nm thick, with protein and carbohydrate biomolecules integrated into it. A variety of lipids such as glycerophospholipid, sphingolipid and sterol, are present in the plasma membrane where they are arranged heterogeneously. At physiological temperatures, the plasma membrane is in liquid-crystalline state. The lipids are mobile both within each monolayer and, to a lesser degree, between monolayers. The plasma membrane contains macrostructures such as rafts (tightly packed lipid regions) and caveolae (small invaginations), and ruffles which may protrude from the external cell surface (Stan 2007). The plasma membrane is selectively permeable. It allows certain substances to cross through diffusion – movement of molecules from a region of higher concentration to a region of lower concentration. Other molecules may enter cells through endocytosis – active enclosure of

Chapter One 3 an extracellular particle within a vesicle bound to the membrane which is in turn transported to intracellular space (Bareford and Swaan 2007). In general, the plasma membrane restricts the passage of therapeutic agents.

1.4 Delivery strategies for therapeutic agents Delivery strategies of anticancer therapeutic agents aim at increasing bioavailability at the intended target to inflict their effect while simultaneously reducing toxic side effects associated with the treatment (Allen and Cullis 2004, Moses et al. 2003, Unger et al. 2004). A number of delivery methods have been developed. Carriers of different size and composition, such as micelles (5-30 nm), liposomes (150-200 nm) and polymers, can be loaded with biologically active molecules with the aim of reducing unwanted interactions of the therapeutic agent with normal tissues (Ferrari 2005, Vegavilla et al. 2008). Delivery of therapeutic agents, incorporated into carriers, to the extracellular space of cancerous tissues is generally improved by exploiting the abnormal tumour morphology and the associated metabolic environment (Jain 2001). Nano-sized carriers (~100- 200 nm) preferentially diffuse from leaky tumour blood vessels versus from the normal tissue vasculature; a phenomenon referred to as enhanced permeability and retention effect (EPR effect) (Besic 2007, Greish 2007). Within the extracellular space of cancerous tissue, molecules of the therapeutic agent can be released from the carrier with the degradation of the carrier body, which can be triggered through the hostile tumour environment. Delivery of therapeutic agents and carriers into the extracellular space of tissue can also be improved using biophysical strategies such as the exposure of ultrasound (Stieger et al. 2007, Visaria et al. 2007). Generally, the biophysical stimulation is applied to the disease site through guidance with imaging systems (Hynynen 2008).

1.4.1 Intracellular delivery strategies Strategies intended to deliver therapeutic agents to the intracellular space aim to improve upon the therapeutic efficacy of molecules, examples of which include genetic and chemotherapeutic agents. Intracellular delivery of therapeutic agents and carriers selectively to diseased tissue can be achieved through endocytosis (Alper 2002). However, this strategy involves identification of receptors that are unique to the disease tissue. The advantage of this approach is that the location of the diseased tissue is not necessary to be identified with imaging systems. Metastatic cancerous lesions, which remain a major challenge, can be treated with this approach. Viral vectors have been developed for gene therapy (Thomas et al. 2003, Xu et al. 2005). However, they suffer from immunogenicity and cytotoxicity.

Chapter One 4 Intracellular delivery can also be achieved using biophysical techniques involving mechanical micro-devices, and ultrasonic and electrical energy. For example, microscopic particles, coated with therapeutic agents, shot at high speeds can penetrate the plasma membrane and deliver the agent locally (Brewster et al. 2006). Although this approach can be efficient, it is not suitable for in vivo applications. The application of ultrasonic or electrical energy has been shown to affect the permeability of cell membranes and to facilitate transportation of cell- impermeable molecules. These phenomena are referred to as sonoporation (Miller et al. 1999) and electroporation (Zaharoff et al. 2008), respectively. Cells exposed to electrical fields demonstrate transient increase in plasma membrane permeability due to the formation of pores across the membrane allowing the passage of extracellular cell-impermeable molecules into cells (Canatella and Prausnitz 2001, Chen et al. 2006). Upon termination of the application of electrical fields, cell membranes recover usually on a time scale of seconds (Teissié et al. 1999). However, electrical pulses cannot be delivered to internal tissues non-invasively without affecting surrounding tissues. In our project, the method chosen to deliver cell-impermeable molecules to the intracellular space of cells is sonoporation. The advantage of ultrasound as a therapeutic system is that it can be focused within the body with millimeter precision to produce specific and localized effects (Chapelon et al. 2000). Ultrasound transducers can be manufactured in any shape and size, and the energy can be directed non-invasively deep within the body. Ultrasound can be applied externally, intravascularly and endoscopically. Tissues and cells located at the transducer focus will be preferentially affected with minimal or no damage to overlying and surrounding tissue structures. Furthermore, the application of therapeutic ultrasound can be guided by a variety of imaging modalities, including ultrasound itself as well as magnetic resonance imaging (MRI) (Hynynen and Clement 2007) to coincide the treatment area with the diseased region. The anatomical and functional information obtained from the imaging system can be used to target and direct the application of therapeutic ultrasound. These characteristics make ultrasound an advantageous system for non-invasive targeted therapeutic-agent delivery applications.

1.5 Ultrasound and microbubbles in imaging and therapy

1.5.1 Physics of ultrasound Ultrasound, sound at frequencies greater than 20kHz, is a mechanical wave that propagates in a medium through the oscillatory motion of particles. In fluids and tissues, ultrasound waves

Chapter One 5 are predominantly longitudinal – the particles are displaced parallel to the propagation direction of the ultrasound wave creating regions of compression (high density) and rarefaction (low density). During propagation, ultrasound waves interact with the underlying tissue resulting in scattering and absorption of the sound energy which have been exploited for imaging and therapeutic purposes, respectively.

1.5.2 Ultrasound in imaging and therapy Ultrasound imaging is a well-established and non-invasive clinical technique that provides information about soft-tissue structure and blood flow in large vessels at rates exceeding 50 frames per second. In medicine, it is used to diagnose diseases through anatomical and functional imaging, guide invasive procedures and monitor tissue response to therapy (Wells 2006). Ultrasound imaging is considered safe and has a long history of clinical use with a long-standing record of safety. However, exposure to ultrasound energy has long been known to produce a variety of biological effects in tissues (Miller et al. 1996). As a therapeutic modality, ultrasound is used in physiotherapy, lithotripsy and more recently in tumour ablation (O’Brien 2007, Yu et al. 2004). The nature of the ultrasound-induced bioeffect is controlled through the exposure parameters. The physical mechanisms involved in ultrasound-induced bioeffects can be thermal and non-thermal. Thermal mechanisms are associated with the absorption of ultrasound energy by the tissue and its conversion into heat. Non-thermal mechanisms are associated with processes that do not produce any significant degree of tissue heating. These non-thermal mechanisms include acoustic cavitation and radiation pressure. Acoustic cavitation is a term that refers to the interaction of acoustic waves with a gas bubble or a gas nucleus. The bubbles can be naturally present (usually in lung and intestine), created through ultrasound exposure (a process known as cavitation inception), or intentionally introduced into the body through injection of microbubble contrast agents (Leighton 2007, O’Brien 2007).

1.5.3 Ultrasound microbubble contrast agents Ultrasound contrast agents consist of encapsulated microbubbles – a shell-encapsulated gas-core bubble. The size distribution of the majority of microbubbles is within 1-5 μm in diameter. The gas core usually consists of perfluorocarbon, nitrogen or sulfur hexafluoride. The shell comprises lipid, albumin-protein or polymer material and its thickness ranges from 10 to 200 nm (Bouakaz and de Jong 2007). Furthermore, the shell structure is not homogeneous; it has a polycrystalline structure (Borden et al. 2005).

Chapter One 6 A microbubble exposed to a pressure wave above a particular amplitude threshold, either oscillates or is disrupted; it is referred to as ultrasound-activated microbubble. The fate and acoustic response of a microbubble depends on a number of factors: (1) its physical properties (e.g. , size, gas and shell), (2) ultrasound exposure parameters, and (3) environmental conditions (e.g., boundary condition, ambient pressure and temperature) (Chen et al. 2003, de Jong et al. 2002).

Stable oscillation A microbubble placed in an ultrasound field above an acoustic threshold will oscillate around an equilibrium radius and emit an acoustic pressure wave. The characteristics of this wave depend on microbubble properties and ultrasound exposure parameters including pulse centre frequency, acoustic pressure and pulse length (Emmer et al. 2007). Microbubbles exposed to sound fields of sufficient amplitude may exhibit a highly non-linear oscillation and emit harmonic, ultra-harmonic and sub-harmonic signals. These nonlinearities have been exploited to detect bubbles which form the basis of contrast-enhanced ultrasound imaging methods (Burns 2002). The maximum oscillation occurs at the resonance frequency of the bubble, which depends on bubble size and shell properties. The resonance frequency of a free gas bubble, assuming a linear and an undamped system, is:

1 3γ P fr = Equation 1.1 2π Ro ρ

where Ro is the initial radius of the bubble, γ is the heat capacity ratio, P is the ambient pressure, and r is the density of the surrounding medium. The behaviour of microbubbles in small vessels is different from unconstrained ones (Caskey et al. 2007). For example, the natural resonant frequency of a 4 μm bubble confined to a compliant vessel with 5 μm inner diameter and 100 μL length increased by a factor of 1.7 compared with a bubble in an unbounded field (Qin and Ferrara 2007). Furthermore, certain bubbles exhibit compression-only behaviour. Phospholipid-coated microbubbles such as Sonovue (Bracco, MI, Italy) are mainly compressed and hardly expanded when exposed to low acoustic pressure (de Jong et al. 2007).

Disruption A unique characteristic of microbubbles is that they can be disrupted by ultrasound. This phenomenon depends on ultrasound exposure parameters. Generally, microbubble disruption

Chapter One 7 increases with acoustic pressure, pulse repetition frequency and decreases with acoustic frequency. As well, the disruption threshold of encapsulated-microbubbles may be higher in smaller vessels (< ~200 μm) (Sassaroli and Hynynen 2007). Microbubbles can be disrupted through various physical mechanisms. The shell of encapsulated microbubbles can be cracked with ultrasound energy releasing the gas and forming a non-encapsulated free gas bubble (Bevan et al. 2008, Postema et al. 2005). These free gas bubbles respond to ultrasound more vigorously than their encapsulated parent microbubbles (Bevan et al. 2007, Wu 2002). Free-gas microbubbles may disappear through passive and ultrasound-driven diffusion. Furthermore, optical observations at high speeds (~10 million frames per second) showed bubbles undergoing fragmentation (Postema et al. 2004). A single microbubble can be fragmented into multiple microbubbles and further respond to ultrasound. Bubbles can also be disrupted through an inertial cavitation mechanism where significant radial expansion of a bubble is followed by rapid and violent collapse. This process produces very high temperatures and pressures, which may damage surrounding tissues through generation of shockwaves and production of free radicals (Atchley et al. 1988). In addition, a liquid microjet may form when the collapse of a bubble occurs near a solid surface, and impinge upon nearby structures (Kodama and Takayama 1998, Ohl et al. 2006).

1.5.4 Ultrasound imaging with microbubbles In imaging, microbubbles are used to improve the detection of small vessels such as in tumours and the myocardium (Wilson and Burns 2006). At the clinical frequencies (1-10 MHz), blood detection in small vessels is challenging due to a combination of factors including low backscattered signal from red blood cells, slow flow, tissue motion, and limited resolution (~ 0.5- 3.0 mm) (Burns 2002, Goertz et al. 2000). In order to increase the amplitude of the signal scattered from blood, ultrasound microbubble contrast agents are administered into the blood compartment (Burns and Wilson 2006, de Jong et al. 2002). A number of ultrasound microbubble contrast agents are currently approved in Canada for diagnostic use including Optison (Nycomed/Amersham, Buckinghamshire, UK), Definity (Lantheus Medical Imaging, Billerica, MA, USA), Albunex (Mallinckrodt Inc, MO, USA) and Sonovue (Bracco, MI, Italy). These agents are considered true intravascular agents – meaning that they remain inside the vasculature. Microbubbles can be targeted to specific disease sites, such as blood clots and endothelial cells (Lindner 2004). The targeting is accomplished with ligands and antibodies, conjugated on the microbubble surface, that bind to cell-receptors associated to the disease such as to VEGF

Chapter One 8 receptors in cancer (Klibanov 2006). More recently, nanobubbles and perfluorocarbon emulsions are being developed for extravascular targeted imaging (Mattrey et al. 1982) and therapy (Liu et al. 2007, Rapoport et al. 2007). This is possible because the smaller size of these particular agents allow them to leak out of the tumour vasculature (Fan et al. 2006).

1.5.5 Ultrasound therapy with microbubbles Microbubble contrast agents found applications in improving therapeutic efficacy of biologically active molecules through enhancement of (1) drug concentration in the vascular compartment of the target area, (2) extravasation through the blood vessels, and (3) intracellular delivery. The application of ultrasound can release molecules of a therapeutic agent loaded within a microbubble. Molecules of the therapeutic agent can be attached to the outer shell, incorporated within the shell and loaded into the interior of microbubbles and released in the vascular compartment through ultrasound-induced microbubble disruption (Ferrara et al. 2007, Unger et al. 2004). Furthermore, the microbubbles can be targeted to the diseased tissue which may further enhance localization (Dayton and Ferrara 2002, Lindner and Kaul 2001, Unger et al. 2001). Extravasation of a therapeutic agent is achieved through permeabilisation of blood vessels with ultrasound and microbubbles (Gaber et al. 1996, Kong et al. 2000). Radiation forces exerted on bubbles displace them in the direction of the traveling sound wave. This force can direct microbubbles loaded with therapeutic agent towards a vessel wall to enhance targeting through receptor-ligand contact (Lum et al. 2006) or to improve extravasation through release of the therapeutic agent close to the vessel wall (Rychak et al. 2007). The application of ultrasound and microbubbles improve intracellular delivery of therapeutic molecules by increasing the permeability of plasma cell membranes (Iwanaga et al. 2007, Kinoshita and Hynynen 2005, Saito et al. 2007, Wei et al. 2004). This phenomenon is termed sonoporation and is the focus of this thesis.

1.6 Sonoporation Sonoporation is a transient and reversible increase in the permeability of plasma cell membranes by ultrasound. This process facilitates transportation of molecules across the cell membrane. Sonoporation has been shown to enhance the intracellular delivery of small compounds (Guzmán et al. 2001), macromolecules (Guzmán et al. 2002, Miller et al. 1999), chemotherapeutic drugs (Yu et al. 2006) and genetic materials (Bao et al. 1997, Kinoshita and Hynynen 2005) in a variety of cell types.

Chapter One 9 One of the earliest investigations on the use of ultrasound to enhance drug delivery was reported by Fellinger and Schmid in 1954, when they enhanced the delivery of hydrocortisone ointment into inflamed tissue (Ng and Liu 2002). In 1995, Tachibana et al. reported that ultrasound in conjunction with microbubbles accelerated thrombolysis, the breakdown of blood clots, with urokinase (a thrombolytic drug) (Tachibana and Tachibana 1995). Sonoporation-mediated drug and gene therapy is currently being developed for cancer, cardiovascular, and blood-brain-barrier- limited treatments (Bekeredjian et al. 2005, Hynynen 2008, Mitragotri 2005). However, it remains a challenge to achieve consistent and controllable delivery, in part because the mechanism of sonoporation is not well understood (Lawrie et al. 2000, Miller et al. 2002, Tachibana and Tachibana 2001). The successful implementation of sonoporation-mediated therapeutic applications depends on understanding its mechanism and the effects of ultrasound and microbubble exposure conditions.

1.6.1 Sonoporation Efficiency Low sonoporation efficiency has limited the development of sonoporation-mediated applications; sonoporation efficiency refers to the percentage of cells which have been reversibly permeabilised. Although the phenomenon of permeabilisation of cell membranes with ultrasound and microubbbles is well established, sonoporation of cells is achieved with large variation in efficiency (Guzmán et al. 2001, Guzmán et al. 2002, Miller et al. 2002, Sundaram et al. 2003). Maximizing sonoporation efficiency depends on increasing the cell membrane permeability while avoiding irreversible membrane damage and cell death. The methods of ultrasound exposure and the experimental conditions vary greatly between studies, limiting the comparison of results (Kinoshita and Hynynen 2007, Miller 2007). Ultrasound exposure parameters including the pulse centre frequency, the acoustic pressure, the pulse duration, the pulse repetition frequency and insonation time are amongst the parameters which differ significantly in sonoporation studies (Figure 1.1). Generally, the ultrasound pulse centre frequency ranged from 0.5 to 5 MHz in sonoporation studies with microbubbles. Some studies suggested that lower ultrasound frequencies (~ 0.5 to 1 MHz) are more efficient at sonoporation (Meijering et al. 2007), whereas other studies suggested the opposite ( ~ 2.5 to 3 MHz). Studies have reported optimum ultrasound insonation time of 30 seconds for endothelial cells and around two minutes for breast cancer (MCF7) cells (Meijering et al. 2007). It has been suggested that molecular uptake and cell viability broadly correlate with transmitted ultrasound energy and that it may serve as a unifying parameter to controlling sonoporation (Zarnitsyn and Prausnitz 2004), however, the dependence on the ultrasound exposure parameters has not been investigated systematically.

Chapter One 10 The microbubble type and concentration also varies in sonoporation studies. Sonoporation of cells can be induced with ultrasound alone. However, in general, sonoporation efficiency is improved when ultrasound treatment is combined with microbubbles (Iwanaga et al. 2007, Kinoshita and Hynynen 2005, Wei et al. 2004). Ultrasound at 500 kHz in the presence of Optison microbubbles were more efficient at gene delivery than low frequency ultrasound at 24 kHz (Zarnitsyn and Prausnitz 2004). Several studies have indicated that the presence and type of microbubbles may be of crucial importance in sonoporation-mediated therapeutic applications (Blomley 2003, Li et al. 2003, Ward et al. 1999). The majority of clinically approved and commercially available microbubble contrast agents have been used in sonoporation studies, including Optison (Kamaev et al. 2004, Miller and Quddus 2000), Definity (Karshafian et al. 2004), Albunex (Miller et al. 1999) and Sonovue (Meijering et al. 2007). Perfluorocarbon gas-filled microbubbles seem more efficient at inducing sonoporation than air-based ones (Blomley 2003). For example, Optison appears to be more effective than Albunex at enhancing gene transfer despite the fact that both agents induced almost identical degrees of mechanical cell death (Li et al. 2003). Sonoporation efficiency is heterogeneous (Guzmán et al. 2001, Guzmán et al. 2002, Miller et al. 2002, Sundaram et al. 2003), which can be partly accounted for by the different experimental conditions including ultrasound dosage and effects associated with experimental apparatus. The variations in ultrasound-mediated gene therapy for different cell lines was shown to be partly attributed to differences in cell membrane properties (Fahnestock et al. 1989). However, sonoporation efficiency varied significantly in cells of the same histological type exposed to the same experimental conditions (van Wamel et al. 2006). Cells are permeabilised, killed or appear to be unaffected by the same treatment. In addition, even among the permeabilised cells, the

Acoustic Pressure

Start Exposure Stop Exposure

Pulse Duration Wavelength = 1/Frequency

Pulse Repetition Period

Figure 1.1: Ultrasound exposure parameters: Pulse centre frequency, peak negative pressure, pulse duration, pulse repetition period and insonation time.

Chapter One 11 number of molecules delivered to the intracellular space varies (Sundaram et al. 2003, Guzmán et al. 2001). The heterogeneous response of cells to sonoporation techniques may be a challenge in sonoporation-mediated applications. A review of the literature indicated that a thorough investigation of the effects of ultrasound and microbubble exposure parameters is still required.

1.6.2 Mechanism of Sonoporation The current conception of the biological mechanism underpinning sonoporation is the formation of non-lethal and transient pores on the surface of cell membranes in a manner that allows cell-impermeable molecules to enter the intracellular space, and subsequent resealing and/or repair of the pores (Mehier-Humbert et al. 2005, Nyborg 2006, Schlicher et al. 2006, Wu 2007). This hypothesis is based on electron microscopy observations of pores formed on plasma membranes of cells exposed to ultrasound and microbubbles. Pores of ~ 2 to 3 μm were observed on ultrasound-treated cells (SK-BR-3, breast cancer cell line) (Zhao et al. 2008). The pores appeared to be transient; they resealed within a few seconds to 24 hours (Zhao et al. 2008). However, the quantification and distribution of these pores have not been well characterized. The acoustical mechanisms underpinning sonoporation include effects associated with stable microbubble oscillation such as microstreaming (Marmottant and Hilgenfeldt 2003, van Wamel et al. 2006), and microbubble disruption by inertial cavitation leading to the generation of shock waves and microjets (Ohl et al. 2006). Microbubbles oscillating in an ultrasound field induce inhomogeneous cyclic pressure fields that create eddy currents in the fluid surrounding the microbubble, which in turn exerts strain on nearby cell membranes (Marmottant and Hilgenfeldt 2003, Wu et al. 2002). Optical measurements showed that cells in the vicinity of oscillating bubbles are deformed (Marmottant et al. 2005). In a study, it was demonstrated that microbubble oscillation deformed the cell membrane locally and only changed its permeability at the deformation site (van Wamel et al. 2006). Bubble disruption does not appear to be necessary for sonoporation. However, processes associated with microbubble disruption appears to contribute to the permeabilisation of cell membranes (Taniyama et al. 2002). Shock waves generated from a collapsing bubble inflict large mechanical stresses on cells in its vicinity. A fluid jet, formed during the asymmetric collapse of microbubbles, may pierce and permeabilise cell membranes (Ohl et al. 2006).

Chapter One 12 1.7 Sonoporation applications The application of sonoporation-mediated therapeutic applications with microbubbles is limited, in principal, to intravascular space: The endothelial cell surface – endothelial cells and junctions – and possibly deeper structures of vessel wall such as smooth muscle cells (Hwang et al. 2005). Cells in the extravascular space may be targeted by direct local injection of microbubbles and therapeutic agent, for example, intramuscularly (Li et al. 2003), or through leakage of perfluorocarbon droplets (Dayton et al. 2006). Perfluorocarbon droplets generally have diameters less than 500 nm and may passively diffuse out of leaky tumour blood vessels (Fan et al. 2006). A number of applications utilizing the bioeffects induced by the combination of ultrasound and microbubbles are being developed. Thrombolysis, the dissolution of blood clots, which is characterized by the deposition of plaques and the consequent reduction of blood flow, is enhanced when thrombolytic agents, such as plasminogen activator (rt-PA), are combined with ultrasound and microbubbles (Meunier et al. 2007, Tsivgoulis and Alexandrov 2007, Zderic et al. 2006). In cancer, sonoporation has been combined with anticancer therapeutic molecules such as bleomycin to enhance anti-tumour cytotoxicity (Larkin et al. 2008, Yu et al. 2006, Yu et al. 2004). Cell death in in vivo tumours was enhanced with the combined treatment in comparison to bleomycin treatment alone. The treatment of brain tumours such as gliomas with anticancer agents is often hindered due to the blood-brain-barrier (BBB). Ultrasound and microbubbles can transiently disrupt the BBB in animals with minimal damage and deliver large molecules to the interstitial space (Hynynen et al. 2001, Meairs and Alonso 2007). MRI-guided focused ultrasound with Optison microbubbles enhanced the therapeutic effect of doxorubicin (DOX), a chemotherapeutic agent, by inducing transient and reversible disruption of the BBB locally (Treat et al. 2007). Furthermore, the treatment of cancer with the combination of anticancer drugs with sonoporation may be more advantageous compared to high-power ultrasound treatment alone. High-power ultrasound therapy can cause sudden metabolic insult and large tumour necrosis, whereas low- intensity ultrasound combined with chemotherapy drugs may induce tumour cell death through apoptosis, which is clinically more desirable (Larkin et al. 2008). However, apoptosis can also be induced with ultrasound alone and in combination with microbubbles (Honda et al. 2002, Vykhodtseva et al. 2006).

Chapter One 13 1.8 Thesis outline The hypothesis guiding this investigation is that sonoporation-mediated intracellular delivery of cell-impermeable molecules is associated with the disruption of the plasma cell membrane, and that these disruptions are induced by ultrasound-activated microbubbles. Experiments were conducted in an in vitro cell suspension system with the aim to improve our understanding of the sonoporation process; the experiments are organized into four chapters (Figure 1.2).

Specific Objectives: a. To determine the effect of ultrasound and microbubble exposure parameters on cell membrane permeability and cell viability. b. To determine the effect of ultrasound exposure parameters on microbubble disruption and establish whether microbubble disruption is necessary in sonoporation. c. To determine if drug size is a limiting factor in sonoporation-mediated therapeutic applications. d. To determine the effect of sonoporation on the long-term viability of cells e. To observe and quantify the size of membrane disruptions generated during sonoporation. f. To determine whether cell cycle affects sonoporation.

In the first study, the effects of ultrasound exposure parameters were investigated (1) on cell membrane permeability and viability by measuring the uptake of cell-impermeable fluorescent molecules, and (2) on microbubble disruption by counting them. Ultrasound parameters including pulse centre frequency, acoustic pressure, pulse duration, pulse repetition frequency and insonation time were varied with the objective to develop insight into the process of sonoporation and to establish an ultrasound parameter space to guide the selection of optimal exposure conditions (Chapter Two). In the second study, the effect of microbubble type and concentration on cell membrane permeability and viability were investigated, and optimal conditions were determined based on maximizing cell permeabilisation. We investigated whether the application of sonoporation is limited (1) by the size of the molecule being delivered (size was varied over two orders of magnitude), and (2) by the ability of cells to proliferate following uptake (Chapter Three). In the third study, the mechanism of sonoporation was investigated through microscopic observations of the plasma membrane and uptake of cell-impermeable molecules. The question of whether the presence of disruptions on the cell membrane (pores) and their size was consistent with the uptake of cell-impermeable molecules was addressed (Chapter Four). In the fourth study,

Chapter One 14 the dependence of sonoporation efficiency on the cell cycle phase was investigated (Chapter Five). Lastly, in Chapter Six, conclusions and preliminary results on sonoporation of in vitro endothelial cells and in vivo blood vessels are reported.

SONOPORATION PHENOMENON: Cell Cell

Cell-impermeable molecule During ultrasound and microbubble Cell-impermeable molecules exposure the cell-impermeable cannot enter the cell molecules cross the cell membrane

(a)

SONOPORATION STUDIES:

Ultrasound Transducer

Ultrasound exposure parameters (Chapter Two)

Mechanism of intracellular delivery Microbubble type microbubble (Chapter Four) and concentration (Chapter Three) Dependence on cell cycle Updake of different (Chapter Five) molecular size markers and clonogenic viability (Chapter Three) (b) Figure 1.2: (a) Sonoporation phenomenon: The application of ultrasound increases the permeability of cell membranes and allows delivery of molecules which othewise would be excluded. (b) Sonoporation studies described in this thesis, which aim at improving our understanding of sonoporation, include the investigation of: (1) the effect of ultrasound exposure conditions, (2) the effect of microbubble exposure conditions, (3) the effect of molecular size on sonoporation-mediated uptake, (4) cell viability, the ability of cells to proliferate, following uptake, (5) the mechanism underpinning sonoporation and (6) the effect of cell cycle phase on sonoporation.

Chapter One 15 CHAPTER TWO

Sonoporation by ultrasound-activated microbubble contrast agents: Effect of acoustic exposure parameters on cell membrane permeability and cell viability

2.0 Abstract This work investigates the effect of ultrasound exposure parameters on the sonoporation of KHT-C cells in suspension by perflutren microbubbles. Variations in insonating acoustic pressure (0.05-3.5 MPa), pulse frequency (0.5-5.0 MHz), pulse repetition frequency (10-3000 Hz), pulse duration (4-32 µs) and insonation time (0.1-900 s) were studied. The number of cells permeabilised to a fluorescent tracer molecule (70 kDa FITC-dextran) and the number of viable cells were measured using flow cytometry. The effect of exposure on the microbubble population was measured using a Coulter Counter. Cell viability and membrane permeability were found to depend strongly on the acoustic exposure conditions. Cell permeability increased and viability decreased with increasing peak negative pressure, pulse repetition frequency, pulse duration and insonation time and with decreasing pulse centre frequency. The highest therapeutic ratio (defined as the ratio of permeabilised to non-viable cells) achieved was 8.8 with 32±4% permeabilisation and 96±1% viability at 570 kPa peak negative pressure, 8 µs pulse duration, 3 kHz pulse repetition frequency, 500 kHz centre frequency and 12 s insonation time with microbubbles at 3.3% volume 2 concentration. These settings correspond to an acoustic energy density (ESPPA) of 3.1 J/cm . Cell permeability and viability did not correlate with bubble disruption. The results indicate that ultrasound exposure parameters can be optimized for therapeutic sonoporation and that bubble disruption is a necessary but insufficient indicator of ultrasound-induced permeabilisation. (This chapter is published in Ultrasound in Medicine & Biology, Vol. 35, No. 5, pp. 847-860, 2009)

Chapter Two 16 2.1 Introduction The quantitative dependence of sonoporation on ultrasound exposure parameters is described in this chapter. We hypothesize that cellular uptake of a high molecular weight marker mediated by ultrasound exposure in the presence of microbubbles can be optimized to achieve high delivery efficiency and minimal cell death. The objectives of this study are to investigate systematically the effect of ultrasound exposure parameters on cell membrane permeability and viability, and to identify optimum ultrasound exposure regimes for cell permeabilisation. In addition, the relationship between ultrasound-induced microbubble disruption and sonoporation efficiency is investigated.

2.2 Methods Cells in suspension were exposed to ultrasound pulses in the presence of Definity microbubbles. Variations in insonating acoustic pressure (0.05-3.5 MPa), pulse frequency (0.5-5.0 MHz), pulse repetition frequency (10-3000 Hz), pulse duration (4-32 μs) and insonation time (0.1- 900 s) were studied. Following exposure of the cells to ultrasound, alterations in cell membrane permeability and cell viability were measured using fluorescent markers and flow cytometry. The disruption of microbubbles was characterized by measuring the size distribution of a sample of microbubbles before and after exposure.

2.2.1 In vitro cell model Cells in suspension were used as the biological model for investigating the sonoporation process. The cells were derived from the murine fibrosarcoma cell line KHT-C (Bristow et al. 1990). They were maintained as monolayers in tissue culture flasks, harvested by trypsinisation, suspended in growth media at a concentration of 1.5x106 cells/mL in a volume of 1.2 mL, and kept at 37°C during the experiment. Prior to ultrasound exposure, microbubbles and cell-impermeable fluorescent molecules were added to the cell suspension, which was adjusted to 1.5 mL volume with phosphate buffered saline (GIBCO 14190, Belbecco’s Phosphate Buffered Saline (1x), Invitrogen, Carlsbad, CA), maintaining a cell concentration of 1.2x106 cells/mL. Following ultrasound exposure, alterations in cell membrane permeability and cell viability were measured using flow cytometry.

Chapter Two 17 2.2.2 Ultrasound exposure system A schematic of the ultrasound exposure apparatus is shown in Figure 2.1. It comprised a single element transducer attached to a micro-positioning system, a waveform generator (AWG520, Tektronix Inc, Beaverton, OR), a power amplifier with receiver circuitry (RPR4000, Ritec Inc, Warwick, RI), a digital acquisition system (Acqiris CC103, Agilent Technologies Inc, Monroe, NY), and a cell exposure chamber with Mylar windows which contained an immersible magnetic stirrer to mix the cells and bubbles. The cell-exposure chamber was of cylindrical shape with 12 mm internal diameter and 10 mm diameter windows across the cylinder with Mylar membranes glued on both sides. The tank was filled with deionized water and maintained at 37°C. The cell suspension was placed at the acoustic focus of the transducer and stirred gently with the magnetic stirrer during the experiment to promote uniform exposure. Three transducers were used in this study: A 500kHz centre-frequency transducer with 32 mm element diameter focused at 85 mm and a –6dB beam width of 31 mm at the focal

Reciever gain:20 dB PC-controlled 12bit A/D System (Acquiris DC440)

Arbitrary Waveform Generator (Tektronix AWG520) Power Amplifier & Receiver (Ritec RPR-4000)

Micropositioner Exposure chamber with acoustic windows Water Heater

Transducer Water Immersible Magnetic Stirrer

Figure 2.1: A schematic diagram of the ultrasound exposure apparatus. Cells are placed within the chamber and exposed to different acoustic conditions.

Chapter Two 18 point (IL0509HP, Valpey Fisher Inc, Hopkinton, MA); a 2.25 MHz centre-frequency transducer with 25.4 mm element diameter focused at 50 mm with 2.8 mm –6dB beam width at the focus (IL0208GP, Valpey Fisher Inc, Hopkinton, MA); and a 5 MHz centre-frequency transducer with 12.7 mm element diameter focused at 25 mm with 1.8 mm –6dB beam width at the focus (IS0504GP, Matec Instruments, Northborough, MA) (Figure 2.2). The –6dB depth-of-field of the three transducers was greater than 20 mm. The transducers were characterized using a calibrated membrane hydrophone (Sonora Medical Systems Inc, Longmont, CO); the spatial peak negative pressure was measured at the focus of the transducer beam in the absence of the exposure chamber. The peak negative pressure was used as the measure of acoustic amplitude. At low pressures, the ultrasonic pulses generated at the focus of the transducer were symmetric, however, the pulses became asymmetric with increasing pressure due to non-linear wave propagation.

2.2.3 Ultrasound microbubble agent A clinically approved and commercially available ultrasound microbubble contrast agent was used: Definity (Perflutren lipid microspheres, Lantheus Medical Imaging, Billerica, MA). Definity microbubbles, activated by shaking the vial for 45 seconds using a Vialmix

(Lantheus Medical Imaging, Billerica, MA), are composed of an octafluoropropane (C3F8) gas core encapsulated within a phospholipid shell. According to the manufacturer, Definity has a concentration of 1.2x1010 bubbles/mL, a mean diameter of 1.1-3.3 μm with 98% less than 10

[kPa] 300 Ultrasound Field 250

200 30 Ultrasound Transducer 150 20 500 kHz 100 10

50 0 0

-10

-20 0

-20 50 -10 Focus 100 0 10 150 20 Figure 2.2: The ultrasound filed of the 500 kHz unfocused transducer with 32 mm element diameter focused at 85 mm.

Chapter Two 19 μm, and a maximum diameter of 20 μm. Before activation, the Definity vial was kept at room temperature for 60 minutes to allow the vial temperature to equilibrate. After activation, the vial was kept for five minutes at room temperature to equilibrate. The vial was resuspended by 10 s of hand agitation and then inverted for 30 seconds before venting with an 18-gauge needle and withdrawing the microbubble suspension into a syringe.

2.2.4 Microbubble size distribution: Coulter Counter The size distributions of Definity microbubbles were measured using a Coulter Counter system (Multisizer III, Beckman Coulter Inc, Fullerton, CA) before and after the microbubbles were exposed to the same acoustic conditions as the cells. A 50 μL volume of Definity microbubbles was added to 1.45 mL of Isoton-II electrolytic solution (Beckman Coulter Inc, Fullerton, CA); this concentration was chosen to match that used in the cell permeability and viability experiments. The Isoton-II solution was filtered twice with 0.1 μm filter before adding the microbubbles. The samples were placed within the exposure chamber, stirred with a magnetic stirrer to maintain a homogeneous solution of bubbles, and exposed to ultrasound. A volume of 1 mL was added to 10 mL volume of filtered isoton-II solution, and the number and size distribution of microbubbles in a 100 μL volume were measured using a 30 μm aperture tube within 60 s of the ultrasound exposure. Measurements were performed for bubble diameters ranging from 0.85 μm to 18 μm. For each sample, three separate aliquots were measured. The number of disrupted microbubbles was calculated over diameters ranging from 1 to 8 μm.

2.2.5 Reversible permeability and PI-viability: PR and VPI Following ultrasound exposure, cell permeability and viability were assessed using fluorescent molecules and flow cytometry. Flow cytometry allows simultaneous measurement of permeability and viability of single cells. A count of 30,000 cells was analysed per sample with flow cytometry (BD FACSCalibur, BD Biosciences International, San Jose, CA). The raw data collected from flow cytometry were processed using FlowJo software (Tree Star Inc., OR). The results are reported as a percent of the 30,000 cells analysed; it is referred to as number of cells.

PI-viability Non-viable cells were identified with flow cytometry (BD FACSCalibur) by the presence of propidium iodide (PI, P-3566, Invitrogen Inc, Carlsbad, CA) within them, an indicator widely used for monitoring cell death, and referred to as PI-dead cells. A 5 μL volume of PI at a concentration

Chapter Two 20 of 0.2 mg/mL was added to a 300 μL volume of the cell suspension within 1.5 hours of ultrasound exposure and just prior to flow cytometry analysis. The number of cells unstained with PI was

counted and referred to as PI-viable (VPI) cells. PI stains cells whose plasma membrane integrity has been compromised; it does not measure the ability of cells to proliferate.

Reversible permeability Alteration in cell membrane permeability was determined with flow cytometry (BD FACSCalibur) by the presence inside the cell of 70 kDa FITC-dextran molecules (fluorescein isothiocyanatedextran, D-1822, Invitrogen Inc, Carlsbad, CA), which normally do not cross intact- cell membranes. A 20 µL volume of the FITC-dextran marker at a concentration of 7.7 mg/mL was added to the cell suspension of 1.5 mL volume 60 s before exposing the cells to ultrasound. Following ultrasound exposure, a volume of 1.2 mL, the remainder of the cell suspension, was rinsed to remove FITC-dextran molecules from the extracellular environment. Next, an anti- fluorescent dye (A-889, Invitrogen Inc, Carlsbad, CA) was added to the cell suspension to quench the fluorescence of FITC-dextran molecules bound to the cell surface. The number of cells stained with FITC-dextran but unstained with PI was counted, indicating reversible cell permeabilisation

(PR). Cells that exhibited fluorescent intensity of FITC-dextran higher than those of the sham control were classified as permeabilised. The fraction of cells whose membrane permeability and PI-viability remained unaltered (unstained with both FITC-dextran and PI molecules) was referred

to as unpermeabilised cells (VUP). The number of VUP cells was determined with Equation 2.1:

= − VVUP PI PR Equation 2.1

2.2.6 Therapeutic Ratio: TRR The “therapeutic ratio” is defined as the number of cells in which permeability was induced

divided by the number of cells that were considered dead by the same exposure. TRR, the ratio of reversibly permeabilised cells to those that are considered non-viable with PI, is defined as:

PR TRR = Equation 2.2 100% − VPI

where PR is the number of reversibly permeabilised cells, and VPI is the number of PI-viable cells. It is a measure of the balance between the desired and destructive effects of ultrasound exposure. The concept of therapeutic ratio was used to identify optimal ultrasound exposure conditions.

Chapter Two 21 2.2.7 Experiments: Ultrasound exposure parameters The effects of the following ultrasound exposure parameters on cell permeability and viability, and on microbubble disruption, were investigated: peak negative pressure, pulse centre frequency, pulse duration, pulse repetition frequency and insonation time. The effect of the peak negative pressure was investigated at different pulse centre frequencies, pulse durations, pulse repetition frequencies and insonation times. Furthermore, the effects of pulse duration and pulse repetition frequency were investigated at constant acoustic energy. The acoustic energy density

(ESPPA) is defined here as the product of spatial-peak-pulse-average intensity (ISPPA) and exposure time (τ):

2 Prms ISPPA = ρ ⋅ c Equation 2.3 τ =PD ⋅ PRF ⋅T

EISPPA =SPPA ⋅τ

where Prms is the root mean square of the peak negative pressure (Pa), r is the density of surrounding medium at 37°C (~1000 kg/m3), c is the speed of sound (~1500 m/s), PD is the pulse duration (s), PRF is the pulse repetition frequency (Hz) and T is the insonation time (s). The exposure time (τ) is defined as the time duration for which the ultrasound is on, whereas the insonation time (T) is defined as the duration of ultrasound treatment from start to finish during which ultrasound is pulsed.

A. Pulse Centre Frequency and Peak Negative Pressure Cells were exposed to three pulse centre frequencies (500kHz, 2MHz and 5MHz) and a range of acoustic pressures, with 32 μs pulse duration, 3 kHz pulse repetition frequency for insonation time of two minutes in the presence of Definity microbubbbles (3.3% v/v). At 500 kHz, the peak negative pressures were 0, 50, 75, 100, 125, 246 and 570 kPa; at 2MHz, the peak negative pressures were 0, 100, 215, 420, 530, 740, 1500 and 2320 kPa; and at 5MHz, the peak negative pressures were 0, 100, 210, 410, 600, 920, 1140, 2440 and 3500 kPa.

B. Pulse Duration and Peak Negative Pressure Cells were exposed to four pulse durations (4, 8, 16 and 32 μs) at each of three acoustic pressures (125, 246 and 570 kPa), with 500 kHz pulse centre frequency and 3 kHz pulse repetition

Chapter Two 22 frequency for two minutes in the presence of Definity microbubbles (3.3% v/v). These settings corresponded to a duty-cycle range of 1.2-9.6%.

C. Pulse Repetition Frequency and Peak Negative Pressure Cells were exposed to five pulse repetition frequencies (10, 100, 200, 1000 and 3000 Hz) at each of three peak negative pressures (125, 246 and 570kPa), with 500 kHz pulse centre frequency and 32 μs pulse duration for two minutes in the presence of Definity microbubbles (3.3% v/v). These settings corresponded to a duty-cycle range of 0.032-9.6%.

D. Constant Acoustic Energy Density (ESPPA) In the previous two sections (B and C), the pulse durations and pulse repetition frequencies were varied, respectively, while the insonation times remained constant at two minutes. In this section, the effects of pulse duration and pulse repetition frequency were investigated by varying each of these parameters with insonation time in order to maintain a constant acoustic energy 2 2 density (ESPPA) of 3.1 J/cm (ISPPA = 10.8 W/cm ), with an exposure time of 288 ms (the total number of cycles transmitted was set to 144,000). For a pulse repetition frequency of 3 kHz, the pulse durations and insonation times were respectively set to the following combinations: 4 μs and 24 s; 8 μs and 12 s; 16 μs and 6 s; and 32 μs and 3 s. For a constant 32 μs pulse duration, the pulse repetition frequencies and insonation times were respectively set to the following combinations: 10 Hz and 900 s; 20 Hz and 180 s; 50 Hz and 180 s; 100 Hz and 90 s; 200 Hz and 45 s; 1 kHz and 9 s; and 3 kHz and 3 s. All the samples were exposed to a peak negative pressure of 570 kPa in the presence of Definity microbubbles (3.3% v/v).

E. Insonation Time Cells were exposed to ultrasound for a range of insonation times at three acoustic pressures (125, 246 and 570 kPa), with 500kHz pulse centre frequency, 32 μs pulse duration and 3 kHz pulse repetition frequency in the presence of Definity microbubbles (3.3% v/v). At 125 kPa, the insonation times were set to 0, 1, 5, 10, 15, 30, 60 and 120 s; at 246 kPa, the insonation times were set to 0, 1, 2, 5, 15, 30, 60 and 120 s; and at 570 kPa, the insonation times were set to 0, 0.1, 0.2, 0.5, 1, 2, 3, 5, 10, 15, 30, 60 and 120 s.

Chapter Two 23 2.3 Results In this chapter, the effect of acoustic pressure, pulse centre frequency, pulse duration, pulse repetition frequency and insonation time - a total of 87 different exposure conditions - on cell permeability, cell viability and microbubble disruption was investigated. The results are shown in Figures 2.3-2.8 as mean and standard error of the mean of at least five independent measurements of cell membrane permeability and cell viability, and three independent measurements of microbubble disruption. In this study, first, the pulse centre frequency that induced maximum permeability was determined with the duty cycle set at 9.6% and insonation time at two minutes. Subsequently, the pulse duration and pulse repetition frequency parameters were varied until permeability and viability reached plateau levels at three different acoustic pressures. The insonation time was varied next and the time where permeability reached a plateau level was identified. Lastly, using the ultrasound settings where permeability was maximized (500 kHz pulse centre frequency, 570kPa peak negative pressure amplitude, 32 μs pulse duration and 3 kHz pulse repetition frequency) and exposure time of 288 ms where permeability approached a plateau level at insonation time of 3 s, cells were exposed to constant acoustic energy density while varying the pulse repetition frequency, pulse duration and insonation time.

2.3.1 Ultrasound exposure parameters

A. Pulse Centre Frequency and Peak Negative Pressure The effects of ultrasound pulse centre frequency and acoustic pressure on cell membrane permeability and cell viability are shown in Figure 2.3. Cell permeability and viability depend on both peak negative pressure and pulse centre frequency parameters when exposed to ultrasound in the presence of microbubbles. Cell viability decreases with peak negative pressure and increases with pulse centre frequency, with a more significant dependence on acoustic pressure at lower frequency (500 kHz). Cell permeability initially increases with peak negative pressure, plateaus and then decreases. Microbubble-mediated ultrasound permeabilisation is also more efficient at lower pulse centre frequencies. In general, higher acoustic pressures and lower pulse centre frequencies are more effective at permeabilising but also at killing cells. Sonoporation is maximized, with 67±2% of the cells permeabilised and 27±4% dead, at 500 kHz frequency and 570 kPa peak negative pressure, which corresponds to an acoustic energy of 125 J/cm2. At this exposure level, the majority of the cells are either permeabilised or dead. At higher pulse centre frequencies, cell

permeability reaches a peak of 51±2% at 2 MHz and 740 kPa peak negative pressure (ESPPA=210

Chapter Two 24 2 2 J/cm ), and a peak of 36±2% at 5 MHz and 2.44 MPa pressure (ESPPA=2.3 kJ/cm ). Above these pressures, cell permeability decreases. Furthermore, at 2 MHz and 5 MHz frequencies, 30-50% of the cells are neither permeabilised nor killed at peak negative pressures of 2.4 MPa and 3.5 MPa, respectively. The effects of ultrasound on size distributions of Definity microbubbles exposed to different acoustic pressures and pulse centre frequencies are shown in Figure 2.4. It is seen that the total number of microbubbles, in the range of 1 to 8 μm in diameter, that remain after ultrasound exposure depends on acoustic pressure and pulse centre frequency (Figure 2.4a). More bubbles are disrupted at lower pulse centre frequencies and higher peak negative pressures. Furthermore, the sizes of bubbles that are disrupted depend on acoustic pressure and pulse centre frequency (Figures 2.4b-d). At 500 kHz frequency and 50 kPa pressure, only microbubbles 4 μm in diameter

Figure 2.3: The effects of peak negative pressure and pulse frequency on cell permeability and viability. The percentages of permeabilised and viable cells are shown with respect to pressure at centre frequencies of 500kHz (solid), 2MHz (dash) and 5MHz (dash-dot), where P is the number of permeabilised cells which are still viable, and V is the number of viable cells. Cell permeability increases and viability decreases with acoustic pressure, and cell permeability decreases and viability increases with centre frequency. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-3.5MPa, f=0.5-5MHz, T=120s, and Definity (3.3% volume concentration).

Chapter Two 25 and larger are disrupted; at 75 kPa peak negative pressure and above, the majority of the bubbles are disrupted (Figure 2.4b). At 2 MHz frequency and 100 kPa pressure, bubbles mainly in the range of 2 to 4 μm in diameter are disrupted; at 420 kPa pressure and above, the majority of the bubbles are disrupted (Figure 2.4c). At 5 MHz frequency and 100-210 kPa pressure, microbubbles smaller than 2.5 μm in diameter are disrupted; at 595 kPa pressure and above all the microbubbles in the range of 1 to 8 μm in diameter are disrupted (Figure 2.4d). This suggested that microbubbles around the resonance frequency, which experience larger volumetric oscilation, were disrupted at lower peak negative pressures.

Figure 2.4: The effects of peak negative pressure and pulse frequency on Definity microbubbles ranging in size between 1 μm and 8 μm. (a) The total number of microbubbles disrupted (per ml) exposed to different peak negative pressures and pulse centre frequencies. The size distribution of Definity (bubbles/ml) exposed to (b) 500 kHz, (c) 2 MHz and (d) 5 MHz for various acoustic pressures. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-2.5MPa, f=0.5-5MHz, T=120s, and Definity (3.3% volume concentration).

Chapter Two 26 B. Pulse Duration and Peak Negative Pressure The effects of pulse duration and peak negative pressure on cell permeability and viability are shown in Figure 2.5. Cell permeability increases and viability decreases with pulse duration, ranging from 4 to 32 μs, at acoustic pressures of 125 to 570 kPa in the presence of microbubbles. It is seen that longer pulses and higher acoustic pressures are more effective at permeabilising cells, but also at killing them. At 32 μs pulse duration and peak negative pressures higher than 250kPa, more than 50% of cells are permeabilised. Increasing the pulse duration results in smaller incremental increases in cell permeability, which tend to plateau; for example, increasing the pulse duration from 4 to 8 μs and from 8 to 16 μs results in ~20% and ~10% increases in cell permeability, respectively. Furthermore, equivalent cell permeabilisation can be achieved at different pulse durations by varying the peak negative pressure; a similar cell permeabilisation of ~40% and cell viability of ~90% can be achieved at an acoustic pressure of 570 kPa and a shorter pulse duration of 4 μs, and at a low acoustic pressure of 246 kPa and a longer pulse duration of 16 μs, which correspond to acoustic energies of 11.6 and 15.6 J/cm2, respectively. The total number of bubbles

Figure 2.5: The effects of pulse duration and peak negative pressure on cell permeability and viability. The percentages of permeabilised and viable cells are shown with

respect to pulse duration, where PR is the number of permeabilised cells which are still

viable, and VPI is the number of viable cells: 125 kPa (solid); 246 kPa (dash); 570 kPa (dash-dot). Cell permeability increases and viability decreases with pulse duration and acoustic pressure. Exposure conditions: PD=0-32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz, T=120s, and Definity (3.3% volume concentration).

Chapter Two 27 in the range of 1 to 8 μm in diameter decreases from ~3x105 to ~1x103 microbubbles per 100 μL volume after exposure to ultrasound at pulse durations of 4 to 32 μs and peak negative pressures of 125 to 570 kPa (data not shown); this corresponds to ~99% bubble disruption.

C. Pulse Repetition Frequency and Peak Negative Pressure The effects of pulse repetition frequency and peak negative pressure on cell permeability and viability are shown in Figure 2.6. Cell permeability increases and viability decreases with pulse repetition frequencies ranging from 10 to 3000 Hz at peak negative pressures of 125 to 570 kPa. Cells exposed to ultrasound at 570 kPa pressure at 10 Hz pulse repetition frequency show 6.2±0.4% permeability and 94±1% viability, whereas at 3 kHz pulse repetition frequency, 67±2% of the cells are permeabilised and 76±4% are viable. Permeability increases almost linearly with pulse repetition frequency at lower acoustic pressures. For pressures of 125 and 246 kPa, these rates of increase were ~8% and ~16%, respectively. However, at 570 kPa pressure, cell permeability plateaus with increasing pulse repetition frequency. In addition, comparable cell permeabilisation

Figure 2.6: The effects of pulse repetition frequency and peak negative pressure on cell permeability and viability. The percentages of permeabilised and viable cells are shown

with respect to pulse repetition frequency, where PR is the number of permeabilised

cells which are still viable, and VPI is the number of viable cells: 125 kPa (solid); 246 kPa (dash); 570 kPa (dash-dot). Cell permeability increases and viability decreases with pulse repetition frequency and acoustic pressure. Exposure conditions: PD=32μs, PRF=10-3000Hz, Pneg=0-570kPa, f=500kHz, T=120s, and Definity (3.3% volume concentration).

Chapter Two 28 of ~25% can be achieved at different pulse repetition frequencies and peak negative pressures: at 100 Hz and 570 kPa, 1 kHz and 246 kPa, and 3 kHz and 125 kPa, which correspond to acoustic energy densities of 6.0, 7.7 and 4.2 J/cm2, respectively. The total number of bubbles in the range of 1 to 8 μm in diameter decreases from ~3x105 to ~2x103 microbubbles per 100 μl volume post- ultrasound exposure at pulse repetition frequencies greater than 100 Hz and acoustic pressures of 125 to 570 kPa (data not shown); this corresponds to ~99% bubble disruption.

D. Constant Acoustic Energy Density (ESPPA) Cell permeability and viability depend on pulse duration and pulse repetition frequency parameters when cells are exposed to ultrasound in the presence of microbubbles (as shown in sections B and C). The effects of pulse duration and pulse repetition frequency on cell membrane 2 permeability and viability at a constant acoustic energy (ESPPA) of 3.1 J/cm are shown in Figure 2.7 with respect to duty cycle (ratio of pulse duration to pulse repetition period in percentage). The 2 exposure time (τ) is set to 288 ms and the intensity (ISPPA) to 10.8 W/cm . These settings are based

Figure 2.7: The effects of pulse duration (dash) and pulse repetition frequency (solid) at 2 constant acoustic energy (ESPPA=3.1 J/cm ) on cell permeability and viability. The percentages of permeabilised and viable cells are shown with respect to duty cycle, where

PR is the number of permeabilised cells which are still viable, and VPI is the number of viable cells. Cell permeability increases and viability decreases with pulse duration and pulse repetition frequency at constant acoustic energy. Exposure conditions: PD=0- 32μs, PRF=10-3000Hz, Pneg=570kPa, f=500kHz, T=3-900s, and Definity (3.3% volume concentration).

Chapter Two 29 on 570 kPa peak negative pressure amplitude, 32 μs pulse duration and 3 kHz pulse repetition frequency, which correspond to the settings where maximum cell permeability was achieved at 500 kHz pulse centre frequency. The 288 ms exposure time is based on 3 s insonation time, 3 kHz pulse repetition frequency and 32 μs pulse duration, which correspond to settings where cell permeability approached a plateau level. It is seen that cell permeability increases and viability decreases with both pulse duration and pulse repetition frequency parameters at constant acoustic energy exposure conditions (Figure 2.7). Cell permeability increases from 4% at 10 Hz pulse repetition frequency to 40% at 3 kHz, for a constant pulse duration of 32 μs. Permeability increases from 25% at 4 μs pulse duration to 40% at 32 μs, for a constant pulse repetition frequency of 3 kHz. At all exposure conditions, ~99% of the microbubbles in the range of 1 to 8 μm in diameter are disrupted (data not shown).

E. Insonation time and Peak Negative Pressure The effects of insonation time and acoustic pressure on cell permeability and viability are shown in Figure 2.8. Cell permeability increases and viability decreases with insonation time at acoustic pressures of 125 to 570 kPa. It is seen that permeability increases with insonation time, then reaches a plateau which depends on acoustic pressure, beyond which permeabilisation is not affected. At 570 kPa pressure, permeability reaches a plateau of ~ 65% after 5 seconds of exposure; at 246 kPa, permeability plateaus at ~60% after 30 seconds; at 125 kPa, permeability plateaus at ~26% after a longer exposure of 60 seconds. A lower degree of permeability is achieved at lower pressures and also at longer insonation times. A similar response is observed for cell viability; at 570 kPa, viability decreases for up to 10 s following the onset of exposure, beyond which viability is not affected.

2.3.2 Optimisation of sonoporation: Therapeutic Ratio

The therapeutic ratios (TRR), the proportion of cells in which reversible permeability increases are induced compared to those that are killed by the same exposure, corresponding to the previous ultrasound exposure parameters are shown in Figure 2.9. This ratio depends on the ultrasound exposure parameters. The therapeutic ratio initially increases with acoustic pressure (Figure 2.9a), pulse duration (Figure 2.9b), pulse repetition frequency (Figure 2.9c), duty-cycle at constant acoustic energy (Figure 2.9d) and insonation time (Figure 2.9e), reaches a maximum and

then decreases. Higher TRRs are achieved at 2 MHz and 5 MHz centre frequencies compared to

500 kHz at the same duty cycle of 9.6% (Figure 2.9a). A TRR of ~6.3 is achieved at 2 MHz and

Chapter Two 30 Figure 2.8: The effects of insonation time and peak negative pressure on cell permeability and viability. The percentages of permeabilised and viable cells are shown with respect

to insonation time, where PR is the number of permeabilised cells which are still viable,

and VPI is the number of viable cells: 125 kPa (solid); 246 kPa (dash); 570 kPa (dash-dot). Cell permeability increases and viability decreases with insonation time and acoustic pressure until they reach a plateau. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-570kPa, f=500kHz, T=0-120s, and Definity (3.3% volume concentration).

420 kPa peak negative pressure and at 5 MHz and 1 MPa acoustic pressure, as compared to a TRR

of 3.2 at 500 kHz and 125 to 246 kPa acoustic pressure. TRRs decrease down to 2.5-3.5 at higher pressure amplitudes for all pulse centre frequencies. The maximum cell permeability of 67% with 76% viability is achieved at 500 kHz frequency, 570 kPa peak negative pressure, 3 kHz PRF, 32

μs PD and 120 s insonation time (Figure 2.3). A peak TRR of ~6 is achieved at 4 μs and 570 kPa,

and at 16 μs and 125 kPa (Figure 2.9b). A peak TRR of 4.4 is achieved at 200 Hz PRF and 570

kPa peak negative pressure (Figure 2.9c). The maximum TRR of 8.8 is achieved at 3 kHz PRF and 8 μs PD (2.4% duty cycle), 500 kHz pulse centre frequency, 570 kPa acoustic pressure and

12 s insonation time (Figure 2.9d). A summary of the ultrasound conditions at maximum TRR and maximum permeability for the settings considered in this study is shown in Table 2.1. To illustrate

the dependence of TRR on acoustic energy density, the data from all the 87 different exposure

conditions is shown in Figure 2.10. The results indicate that there is no correlation between TRR

and energy density. A similar TRR value of ~3.5 can be achieved at acoustic energy densities 2 ranging from 0.5 to 3000 J/cm . The TRR varies from ~0.2 to 8.8 at an acoustic energy density of 3.1 J/cm2.

Chapter Two 31 (a) (b)

Figure 2.9: The therapeutic ratio, the proportion of cells in which reversible permeability increase is induced compared to those that are killed by the same exposure, is shown for various (a) peak negative pressures and pulse frequencies, (b) pulse duration and peak negative pressure, (c) pulse repetition frequency and peak negative pressure, (d) pulse duration and pulse repetition frequency at constant acoustic energy (e) density, and (e) insonation time and peak negative pressure.

Chapter Two 32 Table 3.1: A summary of optimum and maximum therapeutic ultrasound exposure parameters for the conditions considered in this set of studies.

Parameters Highest Therapeutic Maximum Permeability Ratio (8.8) Permeability = 67 ± 2% Permeability = 32 ± 4% Viability = 76 ± 4% Viability = 96 ±1% Pulse frequency 500kHz 500kHz

Peak negative 570kPa 570kPa pressure Pulse duration 8 μs (4 cycles) 32 μs (16 cycles) Pulse repetition 3 kHz 3 kHz frequency Insonation Time 12 seconds 120 seconds

Microbubble Definity (3.3% v/v) Definity (3.3% v/v)

Figure 2.10: The therapeutic ratio, the proportion of cells in which reversible permeability increase is induced compared to those that are killed by the same exposure, is shown 2 as a function of acoustic energy density (J/cm ) for all exposure conditions. TRR values 2 range from ~0.2 to 8.8 at acoustic energy densities of 1-10 J/cm . A similar TRR of ~3.5 can be achieved at energy densities ranging from 0.5 to 3000 J/cm2.

Chapter Two 33 2.3.3 Relationship between permeability and viability To gain insight into the effects of ultrasound exposure on cell membrane permeability and viability, the data from all the exposure conditions are plotted in Figure 2.11. Figure 2.11a shows permeability and viability data as a scatter plot. Figure 2.11b shows a contour plot of therapeutic ratio as a function of permeability and viability. The results show that there is an inverse correlation between cell membrane permeability and viability for all conditions. Cell permeability increases from 0.5% to 70% while viability decreases from 99% to 75%, below which permeabilisation remains constant while viability decreases (Figure 2.11a). Furthermore,

the contour plot identifies regions of high TRR values where permeability approaches ~ 35% and viability decreases to ~95% (Figure 2.11b).

(a)

(b) Figure 2.11: The data for permeabilised cells and viable cells for all exposure conditions are shown in (a). Cell membrane permeability increases to 70% while viability decreases

to 75%. (b) A contour graph of TRR as a function of permeability and viability indicates

areas of high and low TRR values.

Chapter Two 34 2.3.4 Relationship between permeability and microbubble disruption The cell permeability and viability data from all the exposure conditions are plotted with respect to the fractional number of remaining microbubbles in the range of 1 to 8 μm diameter (Figure 2.12). It is seen that there is no correlation between the number of bubbles disrupted and cell membrane permeabilisation or cell death; however, an alteration in cell membrane permeability or cell viability occurred only under ultrasound exposure conditions which resulted in disruption of at least 95% of the bubbles.

Figure 2.12: The number of permeabilised and viable cells with respect to the fractional

number of remaining bubbles is shown here for all exposure conditions, where PR is

the number of permeabilised cells which are still viable, and VPI is the number of PI- viable cells. Bubble disruption appears to be a necessary but insufficient condition to permeabilise cell membranes.

Chapter Two 35 2.4 Discussion This study demonstrates the ability of ultrasound to enhance the uptake of large molecules to which the cell membrane is normally impermeable in an in vitro cell suspension model. The results indicate that sonoporation in KHT-C cells in the presence of microbubbles depends on ultrasound exposure parameters, and that, in this experimental setting, optimisation of sonoporation is possible. This investigation established a parameter space of ultrasound- and microbubble- mediated permeabilisation of cell membranes to guide the selection of optimum exposure conditions, and provides some insight into possible mechanisms of sonoporation.

2.4.1 Optimisation of sonoporation Maximizing the efficiency of intracellular delivery of molecules depends on increasing cell membrane permeability while avoiding irreversible membrane damage or cell death. The

therapeutic ratio (TRR) is defined as one measure of efficacy. This ratio initially increases with acoustic pressure, pulse duration and pulse repetition frequency, reaches a maximum and then decreases. At lower pressures, cell membrane permeability increases and viability decreases with acoustic pressure. However, at higher pressure amplitudes, the number of permeabilised cells decreases. This suggests that high acoustic pressures (> 2-3 MPa) may be less desirable for inducing reversible permeabilisation. Furthermore, lower frequency ultrasound pulses are more effective at permeabilising cells, but also at killing them. Cell permeability in samples treated at lower frequencies (500 kHz) is significantly larger than in those treated at higher frequencies (2-5 MHz). The optimization of pulse duration, pulse repetition frequency and insonation time parameters is conducted at 500 kHz pulse centre frequency, where cell permeability is maximized. Sonoporation efficiency at the higher pulse centre frequencies (2-5 MHz) may be further optimized through pulse duration, pulse repetition frequency and insonation time parameters; this is not addressed in this study. There appears to be a threshold below which cells are neither permeabilised nor killed. This threshold depends on ultrasound pressure, pulse centre frequency and pulse repetition frequency. Peak negative pressure thresholds of 75 kPa, 200 kPa and 600 kPa, at 500 kHz, 2 MHz and 5 MHz ultrasound centre frequencies are observed, corresponding to mechanical indices of 0.11, 0.14 and 0.27, respectively. Above these exposure conditions, approximately 99% of bubbles are disrupted. The Coulter Counter measurements indicate that disruption of Definity microbubbles in the range of 1 to 8 μm in diameter strongly depends on acoustic pressure and pulse frequency but less on pulse duration and pulse repetition frequency for peak negative pressures above 125 kPa and insonation times of a few seconds. The disruption threshold of bubbles increases with pulse centre

Chapter Two 36 frequency. The peaking of therapeutic ratios at 2 MHz and 5 MHz pulse centre frequencies occurs at acoustic pressures that correspond to the minimum pressures required to disrupt the majority of Definity microbubbles. Cell permeability also increases with pulse repetition frequency, pulse duration and insonation time. The results show that there is a dependence on ultrasound pulse duration, with longer pulses generally more efficient at permeabilising as well as killing cells. Guzmán et al. showed that there is no dependence on pulse duration (Guzmán et al. 2001). However, they tested 10-30000 cycles compared to 2-16 cycles per pulse in this study. The results also indicate that pulse repetition frequencies higher than 10 Hz are necessary to obtain significant cell permeabilisation at 500 kHz pulse frequency and 570 kPa peak negative pressure. Although both pulse duration and pulse repetition frequency are significant parameters in ultrasound-induced bioeffects, pulse repetition frequency appears to be a more critical parameter. Ultrasound pulses at high pulse centre frequency, low acoustic pressure and low duty cycle may be optimal for situations where cell survival is important. Low centre frequency, high acoustic pressure and high duty cycle may be better suited when cell survival is unimportant. Customized treatment pulse sequences consisting of various combinations of pulse centre frequencies, acoustic pressures, pulse durations and pulse repetition frequencies may further optimize ultrasound-induced drug uptake. Sonoporation of cells, which is mediated through ultrasound-activated microbubbles, strongly depends on ultrasound exposure parameters; the acoustic behaviour of microbubbles depends on these same conditions. Cells exposed to ultrasound at low pulse repetition frequencies (10 Hz) show no significant alterations in membrane permeability or cell viability. This may be explained by the acoustic behaviour of microbububbles. The shells of microbubbles may be disrupted releasing free gas bubbles, which respond more efficiently to ultrasound and can undergo larger oscillations (Bevan et al. 2007). The median decay of free gas bubbles is on the order of 10- 30 ms, which is consistent with the lack of sonoporation measured at low pulse repetition periods of 100 ms. Alternatively, gas bodies may exist after microbubble disruption and act as cavitation nuclei. The dependence of sonoporation on pulse repetition frequency may also be related to the ability of cells to partially patch or repair their membranes. The literature reports that cells repair their membranes within a few seconds to minutes depending on the exposure conditions (van Wamel et al. 2006, Zhao et al. 2008, Zhou et al. 2008). The question of whether there is a single unifying parameter that can indicate the level of ultrasound-induced bioeffects is important both for safety and optimization of local drug delivery applications. The mechanical index (ratio of peak negative pressure in MPa to square root of the

Chapter Two 37 pulse centre frquency in MHz) is not a good indicator of ultrasound and microbubble induced bioeffects in our experimental set up. Cells exposed to the same mechanical index and duty cycle show a higher permeabilisation at lower pulse centre frequencies, but also lower viability. At an MI of 0.55, 58% of cells are permeabilised at 500 kHz frequency, compared with 50% at 2 MHz and 30% at 5 MHz. Studies have suggested that molecular uptake, cell viability, and cell transfection broadly correlate with ultrasound energy exposure (Zarnitsyn and Prausnitz 2004). They have shown that all the data can be roughly incorporated into a single curve with optimum energy 2 exposures in the range of 10 to 40 J/cm . Our results, however, show that acoustic energy (ESPPA) is not a unifying parameter that can predict both cell membrane permeabilisation and cell death at all exposure conditions as shown in Figure 2.10. Rationally, acoustic energy (ESPPA), which is defined as the product of intensity (ISPPA) and exposure time (τ), cannot be a predictive parameter of ultrasound-induced bioeffects in the presence of microbubbles. Sonoporation depends on pulse centre frequency, which is not accounted for in the acoustic energy calculations. It is true that under certain exposure conditions, acoustic energy can be a predictor of optimum exposure conditions for sonoporation (e.g. in the range of 2-20 J/cm2). However, cells exposed to constant acoustic energy still showed a dependence on pulse duration and pulse repetition frequency. Furthermore, cell permeability and viability plateau with insonation time, indicating that further exposure does not cause any additional permeabilisation. It is not surprising that sonoporation cannot be predicted by a single parameter. The acoustic response of microbubbles is highly non-linear and strongly depends on the ultrasound exposure parameters (de Jong et al. 2007, Emmer et al. 2007). Real-time monitoring of microbubble activity during treatment with ultrasound may be necessary to predict and control the sonoporation process.

2.4.2 Mechanism of sonoporation Bubble disruption is a phenomenon by which ultrasound energy can be concentrated into a small volume. The energy released during bubble disruption can create transient pores in cell membranes allowing macromolecules to pass; it can also cause cell lysis (Taniyama et al. 2002). Figure 2.11 shows an inverse correlation between cell permeabilisation and cell viability, suggesting that a single process might well be responsible for both permeabilising and killing cells. Sonoporation may cause membrane damage in all cells, with some cells able to repair themselves to remain viable, while others are fatally damaged by irreversible membrane permeabilisation. If this is the case, sonoporation could be considered as transitory damage to surviving cells.

Chapter Two 38 2.4.3 Role of microbubble disruption The Coulter Counter measurements do not reflect the complex acoustical history of the microbubbles: they can undergo stable oscillation, disruption, collapse, fragmentation and coalescence, or a combination of these, depending on the ultrasound exposure conditions and history (Bevan et al. 2007, Church and Carstensen 2001). Substantial cell permeabilisation (more than 5%) was associated with bubble disruption, as shown by the Coulter Counter (Figure 2.12). Bubble disruption is seen to be a necessary but not a sufficient condition to change cell membrane permeability or to kill cells. The results do not imply that the mechanism of sonoporation is bubble disruption. A detailed theory to predict the extent of ultrasound and microbubble-mediated bioeffects is not available (Miller et al. 2008), though a number of different mechanisms may be responsible. The physical processes associated with microbubble collapse (inertial cavitation) in the vicinity of the bubble include generation of high shear stresses, high local temperatures, shock waves and free radicals. Bubbles in the vicinity of cells undergo asymmetrical collapse, which can result in high-speed fluid microjets (Ohl et al. 2006). These processes usually increase with ultrasound intensity and decrease with frequency. The acoustic response of microbubbles during exposure and the rate of microbubble disruption can provide additional information which may elucidate the mechanisms that mediate the transmembrane transport of extracellular molecules into cells. In this study, the ultrasound exposure conditions can be considered free of thermal bioeffects. However, microbubbles can significantly reduce the ultrasound threshold to induce thermal damage (Razansky et al. 2006).

2.4.4 Limitations of the study There are several limitations to the present study. A variety of in vitro cell exposure systems have been utilized for studying ultrasound-induced bioeffects in cells, including exposure through static and rotating test-tubes (glass/plastic) (Miller et al. 1996). Cell suspension systems allow considerable control of environmental, biological and exposure parameters including cell density and microbubble concentration. However, careful standardization of both the acoustical and biological procedures is required to improve comparisons between studies and also to increase the relevance of in vitro studies to in vivo ones. In vitro sonoporation experiments should be carried out in acoustically favourable conditions; acoustically semi-transparent tubes and dishes, which distort the ultrasound beam characteristics, should be avoided. In addition, standing waves, which may be generated due to internal reflections in plastic tubes, should be taken into consideration. Furthermore, the effect of radiation force, which acts in the direction of the wave proparagion,

Chapter Two 39 on microbubbles should be taken into consideration. This force can displace the microbubbles to the rear of the exposure chamber resulting in a non-uniform mixing of bubbles and cells. The dimensions of the cell-exposure chamber and the ultrasound beam characteristics should be specified. The cells should be maintained at biologically relevant temperature throughout the experiment at 37°C. Furthermore, the handling of microbubbles, which may be significant, should also be standardized. In this study, the acoustic pressure, which was measured in the absence of the exposure chamber, can be influenced by the chamber, the Mylar windows, microbubble attenuation and non-linear propagation. Furthermore, the beam-characteristics at the focus can change the expected resident time of cells within the ultrasound beam. For the 2 MHz and 5 MHz transducers, the effective resident time of cells is approximately 14 s and 4 s, respectively, for two minutes of ultrasound exposure duration. Cell permeability and viability were measured two hours after treatment. Cell viability was measured with propidium iodide, which does not measure the ability of cells to proliferate. Permeability was measured using a 70 kDa FITC-dextran marker: the dependence on different size markers was not addressed here. In addition, cells were exposed to ultrasound in the presence of only one type of microbubble (Definity agent) and at only one concentration (3.3% v/v). Based on the concentration of Definity specified by the manufacturer (Lantheus Medical Imaging, Billerica, MA), which is 1.2×1010 after activation of the vial, there is about 4×108 bubbles/ml in a sample. This corresponds to ~330 bubbles per cell. However, Definity contains a large number of small microbubbles below 1 μm diameter. Here, the number of microbubbles in a sample within 1-8 μm diameter range, measured with Coulter Counter, is ~ 33×106 bubbles/mL; this corresponds to ~27 bubbles per cell. The dose used in this study was ~ 2.8 ml/kg based on a 60 kg person containing 5 L of blood at 3.3% microbubble volume concentration. This concentration exceeds the recommended human dose of 10-30 μLkg. However, dosing studies in primates indicate that 10 mL/kg of Definity can be safely administered without adversely affecting hemodynamics or blood pressure (Unger et al. 2004). The acoustic behaviour of the microbubbles was not measured in this study. Future work will further investigate the mechanism of sonoporation with passive detection of cavitation and acoustic attenuation measurements through a population of microbubbles. In this study, KHT-C cells are used as a biological model for investigating the sonoporation process, and not as an in vitro model of an in vivo system. Whether the general conclusions of this in vitro study may be generalized to other cell types has yet to be investigated. In addition, the observed dependence of ultrasound-induced bioeffects on the exposure parameters in an in vitro

Chapter Two 40 cell suspension model is likely to differ from in vivo systems, where cells are integrated together and with the extracellular matrix to form a tissue. These differences need to be addressed by future studies.

2.5 Conclusions Sonoporation of KHT-C cells in suspension with microbubbles was observed for a range of ultrasound exposure conditions. It was found that ultrasound causes either reversible cell membrane permeabilisation which in turn allows cell membrane-impermeable molecules to pass, or it causes irreversible cell membrane damage that results in cell death. Cell permeability and non-viability increase with acoustic pressure, duty cycle and insonation time and decrease with ultrasound frequency. Sonoporation was maximized with 67±2% permeabilisation and 24±4% non-viability at 570kPa peak negative pressure, 32 μs pulse duration, 3 kHz pulse repetition frequency, and 120 s insonation time with Definity microbubbles at 3.3% volume concentration. Optimal ultrasound conditions were identified using the concept of a therapeutic ratio, defined as

the ratio of permeabilised to non-viable cells. A maximum TRR of 8.8 was attained with 32±4% permeabilisation and 96±1% viability at 570 kPa peak negative pressure, 8 μs pulse duration, 3 kHz pulse repetition frequency, 500 kHz centre frequency and 12 s insonation time with Definity microbubbles at 3.3% volume concentration. Cell membrane permeabilisation was associated with bubble disruption, however, bubble disruption was determined to be a necessary but insufficient indicator of sonoporation. Additional work is needed to examine other cell lines and identify comparable conditions in vivo.

Chapter Two 41 CHAPTER THREE

Microbubble mediated sonoporation of cells: Clonogenic viability and influence of molecular size on uptake

3.0 Abstract This work investigates whether the application of sonoporation is limited by the size of a macromolecule being delivered and by the ability of cells to proliferate following uptake. KHT-C cells in suspension were exposed to variations in ultrasound pressure (0 to 570 kPa) and microbubble shell-type (lipid and protein) at fixed settings of 500 kHz centre frequency, 32 μs pulse duration, 3 kHz pulse repetition frequency and two minutes insonation. Reversible permeability (PR), defined as the number of cells stained with FITC-dextran and unstained with propidium iodide (i.e., PI- viable), was measured with flow cytometry for marker molecules ranging from 10 kDa to 2 MDa in size. Viable permeability (PV) defined as the number of permeabilised cells that maintained their ability to proliferate, was measured by clonogenic assay. Comparable intracellular delivery of all sizes of molecules was achieved, indicating that intracellular delivery of common therapeutic drugs may not be limited by molecular size. Maximum PR’s of 80% (at 10 kDa) and 55% (at 10 kDa) were achieved with lipid coated bubbles at 3.3% v/v and protein coated bubbles at 6.7% v/v concentrations. The PI-viability was approximately 80% at 570 kPa in both cases. The maximum

PV achieved with both agents was 22%, while inducing a lower overall clonogenic viability with the lipid (39%) compared to the protein (56%) shelled bubbles. This study demonstrates that large macromolecules, up to 2 MDa in size, can be delivered with high efficiency to cells which undergo reversible permeabilisation, maintaining long-term viability in approximately half of the cells. (This chapter was submitted to Ultrasonics)

Chapter Three 42 3.1 Introduction This study investigates whether the efficacy of sonoporation is limited by the size of a macromolecule being delivered and by the ability of cells to proliferate following uptake. In Chapter Two, it was established that sonoporation depends on ultrasound exposure parameters, and that 70 kDa FITC-dextran molecules can be delivered at high efficiency when present during treatment; reversible permeabilisation of ~70% was achieved. Reversibility of membrane permeabilisation was assessed by the exclusion of PI molecules; PI was added 90 minutes after termination of the ultrasound exposure. PI stains cells whose plasma membrane integrity has been compromised but does not measure the ability of intact cells to proliferate. Long-term viability of cells following sonoporation is clearly important. This study examines a range of molecular sizes, over two-orders of magnitude, under similar exposure conditions and thus addresses whether sonoporation efficiency is dependent on the size of the macromolecule being delivered. Furthermore, it measures the ability of cells which have been permeabilised at high efficiency to proliferate using a clonogenic assay. The hypothesis guiding this study is that sonoporation mediates similar intracellular delivery of molecules with different molecular sizes and that reversibly permeabilised cells are viable and able to proliferate. The objectives of this study are to determine the optimum microbubble (type and concentration) based on maximizing cell permeability, to measure intracellular uptake of different molecular weight FITC-dextran markers ranging from 10 kDa to 2 MDa in size, and to measure the clonogenic viability of permeabilised cells.

3.2 Methods Cells in suspension were exposed to ultrasound and microbubbles. Variations in microbubble type and concentration, acoustic pressure and permeability marker size were studied. The effects on cell membrane permeability and viability were measured with flow cytometry and clonogenic assay.

3.2.1 In vitro cell model KHT-C cells in suspension were used in this study (Section 2.2.1).

3.2.2 Ultrasound exposure system The ultrasound exposure apparatus and the 500 kHz centre frequency transducer was used to expose cells to ultrasound (Section 2.2.2).

Chapter Three 43 3.2.3 Ultrasound microbubble agent Two clinically approved and commercially available ultrasound microbubble contrast agents were used: Definity and Optison (Perflutren protein microspheres, Nycomed/Amersham, Buckinghamshire, UK). The activation and handling of Definity microbubbles is outlined in

Section 2.2.3. Optison microbubbles are composed of an octafluoropropane (C3F8) gas core encapsulated within a human serum albumin shell. According to the manufacturer, Optison has a concentration of 5-8x108 bubbles/mL, a mean diameter of 3-4.5 μm with 95% less than 10 μm. Optison microbubbles were prepared by gently rotating the vial for 30 seconds until it was completely resuspended. After activation, the vial was vented with an 18-gauge needle and the bubble suspension was withdrawn into a syringe.

3.2.4 Reversible permeability and PI-viability: PR and VPI Reversible permeability and PI-viability were measured using fluorescent molecules and flow cytometry. Non-viable cells were identified by the presence of PI molecules within them (Section 2.2.5). Reversible permeability was assessed using four sizes of FITC-dextran molecules: 10 kDa (D-1820, 43.1 mg/mL), 70 kDa (D-1822, 7.7mg/mL), 500 kDa (D-7136, 2.5mg/mL) and 2 MDa (D-7137, 2.5mg/mL) (Invitrogen Inc, Carlsbad, CA). FITC-dextran was used as permeability tracer since it is available in a wide range of molecular weights. The protocol is specified in Section 2.2.5.

3.2.5 Clonogenic Assay: CRP and CUP The long-term viability of sonoporated cells was measured through their ability to proliferate and form a cell colony. Following ultrasound exposure and using a FacsDiva sorter (BD Biosciences International, San Jose, CA), cells were sorted into two subpopulations: Reversibly permeabilised (PR) and unpermeabilised (VUP). Cells were plated into Petri dishes and after four days of incubation stained with methylene blue (1% w/v, VWR International, Ontario, Canada). Cells retaining their reproductive ability formed colonies, a cluster of at least 15 cells, and were considered clonogenically viable. The number of cell colonies formed was counted manually and normalized with respect to the control. Cell clonogenicity of the two subpopulations is referred to as CRP for the reversibly permeabilised cells, and CUP for the unpermeabilised cells.

Chapter Three 44 3.2.6 Experiments Three experiments were conducted. First, the effect of the concentration of the two microbubble agents (Definity and Optison) on cell membrane permeability and viability was investigated and optimal concentrations based on maximizing cell permeability identified. Next, intracellular delivery of different sized molecules was measured with both agents at their optimal concentrations over a range of acoustic pressures. Last, the clonogenicity of sonoporated cells was measured. The ultrasound exposure conditions were set to 500 kHz pulse centre frequency, 32 µs pulse duration, 3 kHz pulse repetition frequency, and two minutes insonation time. The peak negative pressure was varied from 0 to 570 kPa. The experiments were repeated at least four times and the results were reported as the mean value together with standard error of the mean.

A. Microbubble Agent Cells were exposed to ultrasound pulses at six concentrations of each of the two agents. A bubble suspension volume of 0, 1, 10, 50, 100 or 200 µL was added to the cell suspension and adjusted to 1.5 ml volume with phosphate buffered saline (GIBCO 14190, Dulbecco’s Phosphate Buffered Saline (1x), Invitrogen, Carlsbad, CA), which corresponded to 0, 0.067%, 0.67%, 3.3%, 6.67% and 13.2% volume concentration (v/v). The ultrasound peak negative pressure was set at an amplitude of 570 kPa, based on previously established optimal ultrasound exposure conditions (Karshafian et al. 2009) (Chapter Two). Reversible permeability (with 70 kDa FITC-dextran) and PI viability were measured using flow cytometry. The optimal microbubble concentrations of the two agents that maximized reversible permeabilisation were determined. Each measurement was repeated four times.

B. Uptake of Different Molecular Size Markers Cells were exposed to ultrasound pressures of 125, 246 and 570 kPa with each of the two agents (Definity and Optison) at their optimal concentrations. Cell permeability to FITC-dextran molecules of 10 kDa, 70 kDa, 500 kDa and 2 MDa and PI viability were measured at each of the seven exposure settings (one control and six treated) using flow cytometry. Each measurement was repeated eight times at each of the seven exposure conditions with each of the four FITC- dextran markers. Confocal fluorescent microscopy (Zeiss LSM510, Carl Zeiss MicroImaging Inc, Thornwood, New York, USA) images were acquired of sample cells stained with 10 kDa FITC- dextran following exposure to ultrasound and Definity microbubbles.

Chapter Three 45 C. Clonogenic Viability of Permeabilised Cells Cells were exposed to the three ultrasound pressures above with each of the two agents.

Following exposure, cell clonogenicity was measured for reversibly permeabilised (CRP) and unpermeablised (CUP) cells at each of the six exposure conditions. For each of the 14 cell groups, 1×103 cells from four independent samples were plated in triplicate. The number of cell colonies formed after four days of incubation were counted and normalized with respect to the control.

Flow cytometry measured the number of PI-viable (VPI) and reversibly permeabilised (PR) cells. Clonogenic assay determined the clonogenicity of permeabilised (CRP) and unpermeabilised

(CUP) cells with the same number of cells plated. The number of the reversibly permeabilised (PR) and the PI-viable (VPI) cells that remained clonogenically viable were determined by multiplying the percentage obtained with flow cytometry by its corresponding clonogenicity level. The number of cells which are permeabilised and clonogenically viable (PCV), and the number of clonogenically viable cells (VC), are thus:

PP= ⋅C CV RRP Equation 3.1

VPCR= ⋅CVRP +UP ⋅CUP

where PR is the number of reversibly permeabilised cells, CRP is the clonogenicity of reversibly permeabilised cells, VUP is the number of unpermeabilised cells, CUP is the clonogenicity of unpermeabilised cells, and VPI is the number of PI-viable cells. The number of reversibly permeabilised cells (PR) obtained with the 70 kDa FITC-dextran marker was used to match the molecular size of the FITC-dextran in the clonogenic assay measurement.

3.2.7 Therapeutic Ratio: TRR and TRC Two therapeutic ratios were calculated: One based on reversible permeability and PI-dead cells (TRR), and the other on viable permeability and clonogenically dead cells (TRC). TRR is defined in Equation 2.2. TRC is defined in Equation 3.2, as:

PCV TRC = Equation 3.2 100% − VC

where PCV is the number of clonogenically viable and permeable cells, and VC is the number of clonogenically viable cells.

Chapter Three 46 3.3 Results Ultrasound and microbubbles induced reversible permeabilisation in 80% of the cell population (24,000 of the 30,000 cells analysed with flow cytometry were stained with 10 kDa FITC-dextran and unstained with PI molecules) (Figure 3.1). A scatter diagram of the fluorescence from the FITC-dextran marker (10 kDa) against PI is shown in Figure 3.1a. The upper-left hand quadrant contains cells that were reversibly permeabilised. Intracellular delivery was quantified among the cells remaining PI-viable. Confocal microscopy (Figure 3.1b) of a cell exposed to ultrasound and microbubbles shows a 1 µm slice, confirming that the FITC-dextran is within the cell.

3.3.1 Microbubble agent The effect of microbubble concentration on permeability and PI-viability with Definity and Optison agents are shown in Figures 3.2 and 3.3, respectively. Ultrasound alone induced ~2% increase in cell permeability while PI-viability remained unaffected. With increasing microbubble concentration, PI-viability decreased and cell permeability increased reaching a maximum, beyond which it decreased. Reversible cell permeability reached a maximum of 71±2% (70 kDa FITC- dextran) and PI-viability of 79±1% with Definity at 3.3% (v/v) (Figure 3.2). Optison achieved a maximum reversible permeability of 44±1% (70 kDa FITC-dextran) with PI-viability of 82±1% at 6.7% (v/v) (Figure 3.3).

4 Viable Nonviable 10 3 Uptake 10 2 10 1 Uptake No FITC-dextran Intensity 10 0 10 100 101 102 103 104 (a) Propidum Iodide Intensity (b) Figure 3.1: Sonoporation of KHT-C cells with 10 kDa FITC-dextran. (a) Flow cytometry fluorescence intensity of FITC-dextran versus propidium iodide (PI). The upper-left hand quadrant shows cells which are permeabilised and viable. Positive PI staining indicates cell death. At these conditions, 78±2% of the cells were permeabilised, with only 18±1.5% killed. (b) Fluorescent confocal microscopy of a permeabilised cell of 1µm thickness showed the 10 kDa FITC-dextran molecules inside the cell. Exposure conditions: PD=32µs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=120s, and Definity (3.3% v/v).

Chapter Three 47 Figure 3.2: Cell permeability (PR) and PI-viability (VPI) with Definity agent; microbubble

concentration was varied from 0 to 13.2% v/v. PR, cells stained with 70 kDa FITC- dextran and unstained with PI (PI-viable) molecules, increased with microbubble concentration and reached a maximum level of 71±2% % at 3.3% v/v, beyond which it decreased. The 3.3% v/v was identified as the optimal concentration of Definity.

VPI decreased with microbubble concentration. Exposure conditions: PD=32µs, PRF=3kHz, Pneg=570kPa, f=500kHz and T=120s.

Figure 3.3: Cell permeability (PR) and PI-viability (VPI) with Optison agent; microbubble

concentration was varied from 0 to 13.2% v/v. PR, cells stained with 70 kDa FITC-dextran and unstained with PI (PI-viable) molecules, increased with microbubble concentration and reached a maximum level of 44±1% % at 6.7% v/v, beyond which it decreased. The

6.7% v/v was identified as the optimal concentration of Optison. VPI decreased with Optison microbubble concentration. Exposure conditions: PD=32µs, PRF=3kHz, Pneg=570kPa, f=500kHz and T=120s.

Chapter Three 48 3.3.2 Uptake of different molecular size markers Permeabilisation and PI-viability of cells to molecules ranging of different size are shown in Figure 3.4 (using Definity at 3.3% v/v) and Figure 3.5 (using Optison at 6.7% v/v). Comparable intracellular delivery of all the markers (10 kDa, 70 kDa, 500 kDa and 2 MDa) was achieved at this exposure condition. This result was independent of microbubble type and acoustic pressure. However, the number of reversibly permeabilised cells depended on microbubble type. Maximum permeability of 80% was achieved with Definity (Figure 3.4). With Optison, a maximum permeability of 55% was achieved at 570 kPa pressure (Figure 3.5). Furthermore, reversible permeability increased with acoustic pressure. PR increased threefold from 20-30% to 60-70% with an increase in pressure from 125 kPa to 246 kPa using Definity (Figure 3.4).

Figure 3.4: Uptake of different molecular weight FITC-dextran markers with Definity agent. Comparable uptake of the FITC-dextran marker of 10 kDa, 70 kDa, 500 kDa and 2 MDa in size was achieved at fixed acoustic pressure. Reversible permeability

(PR) increased with acoustic pressure for all marker sizes. Maximum PR of 78±2% was achieved with 10 kDa FITC-dextran at 570 kPa. Exposure conditions: Definity (3.3% v/v), PD=32µs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s.

Chapter Three 49 Figure 3.5: Uptake of different molecular weight FITC-dextran markers with Optison agent. Comparable uptake of the FITC-dextran marker of 10 kDa, 70 kDa, 500 kDa and 2 MDa in size was achieved at fixed acoustic pressure. Reversible permeability

(PR) increased with acoustic pressure for all marker sizes. Maximum PR of 56±4% was achieved with 10 kDa FITC-dextran at 570 kPa. Exposure conditions: Optison (6.7% v/v), PD=32µs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s.

3.3.3 Viability of reversibly permeabilised cells The percent of cell colonies formed by reversibly permeabilised and unpermeabilised cells following exposure is shown in Figure 3.6. These percentages are normalized with respect to the untreated control; the corresponding numbers of clonogenically viable (VC) and permeable

(PVC) cells are shown in Figure 3.7. Among the cells exposed to ultrasound and microbubbles, clonogenicity of reversibly permeabilised cells (CRP) was less than the unpermeabilised ones (CUP) at all acoustic pressures and with both agents (Figure 3.6). Clonogenicity of reversibly permeabilised cells with Definity and Optison microbubbles decreased to 30-38% and 48-51%, respectively. In contrast, unpermeabilised cells exposed to ultrasound with Definity or Optison maintained

Chapter Three 50 their colony forming ability in 70-98% and 89-91% of the cells, respectively. A maximum viable permeability of 22±2% with 39±2% clonogenic viability was achieved with Definity (Figure 3.7). With Optison, viable permeability was maximized at 21±1% with 56±1% clonogenic viability (Figure 3.7). A comparable viable permeability of ~22% was achieved with both agents, however, a much lower clonogenic viability was found with Definity (VC = 39%) compared to Optison (VC = 56%).

Figure 3.6: Clonogenicity of sonoporated cells. The number of cell colonies formed by reversibly permeabilised and unpermeabilised cells, normalized with respect to untreated control, is shown against acoustic pressure for two microbubble agents (Definity at 3.3% v/v and Optison at 6.7% v/v). Clonogenicity of reversibly permeabilised cells with

Definity (D-CRP) and Optison (O-CRP) decreased to 30-38% and 48-51%, respectively. Clonogenicity of unpermeabilised cells following ultrasound and microbubble treatment

remained high; clonogenicity of unpermeabilised cells with Definity (D-CUP) and

Optison (O-CUP) was 70-98% and 89-91%, respectively. Clonogenicity of reversibly permeabilised cells was lower compared to clonogenicity of unpermeabilised cells. Exposure conditions: PD=32µs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s.

Chapter Three 51

Figure 3.7: Permeabilised and clonogenically viable cells. Viable permeability (PCV), the number of cells which are permeabilised and clonogenically viable, and clonogenic

viability (VC), the number of clonogenically viability (VC), are shown with two agents (Definity and Optison) and four acoustic pressures (0-570 kPa). Maximum viable

permeability of 22% was achieved with both agents; D-PCV and O-PCV refer to viable permeability with Definity and Optison, respectively. A higher clonogenic viability was

achieved with Optison compared to Definity; D-VC and O-VC refer to clonogenic viability with Definity and Optison, respectively. Exposure conditions: PD=32µs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s.

Chapter Three 52 3.3.4 Therapeutic Ratio: TRR and TRC

Therapeutic ratios of reversible (TRR) and viable (TRC) permeabilisation of KHT-C cells

are shown in Figure 3.8. TRR and TRC, the proportion of cells in which permeability increases are induced compared to those that are killed by the same exposure, were determined based on reversible

permeability (Figures 3.4 and 3.5) and viable permeability (Figure 3.7) data, respectively. A TRR of 3.4±0.1 was achieved with Definity compared to a TRR of 2.4±0.1 with Optison. However, a higher TRC was achieved with the Optison agent. A TRC of 0.48±0.03 was achieved with Optison compared to 0.36±0.03 with Definity microbubbles. The therapeutic ratios depended on acoustic pressure and microbubble type.

Figure 3.8: Therapeutic ratio of sonoporated cells with viability assessed using

propidium iodide (PI) and clonogenic assay. The therapeutic ratios, TRR and TRC, the number of permeabilised cells divided by the number that are killed by the same exposure, are shown with two agents (Definity and Optison) and four acoustic pressures (0-570

kPa). A higher TRR, viability assessed with PI, was achieved with Definity (D-TRR) than

Optison (O-TRR). However, a higher TRC, viability assessed with colony assay, was

achieved with Optison (O-TRC) than Definity (D-TRC). Exposure conditions: PD=32µs, PRF=3kHz, Pneg=0-570kPa, f=500kHz and T=120s.

Chapter Three 53 3.4 Discussion The molecular weight cut-off for molecules deliverable to permeabilised cells by ultrasound and microbubbles, which subsequently resealed, is larger than 2 MDa in size. This implies that molecular size may not be a limiting factor in sonoporation-mediated therapies. Guzman et al. have shown similar intracellular uptake for small (623 Da) and intermediate size (464 kDa) molecules (Guzmán et al. 2002). However, sonoporation induced by lithotripter shock waves in in vitro cells showed a dependence on molecular size (Gambihler et al. 1994). This study suggests that pores as large as 56 nm, which correspond to the diameter of 2 MDa molecules, are formed on the membrane of cells exposed to ultrasound and microbubbles which allows molecules to enter the intracellular space. Following intracellular delivery, cells reseal their membrane as indicated by the exclusion of PI marker molecules.

3.4.1 Reversible and viable permeabilisation This study demonstrated that reversibly permeabilised cells were viable and able to proliferate. The long-term viability of sonoporated cells is an important consideration for potential applications in drug delivery. Propidium iodide, which is commonly used in sonoporation studies, may not be a reliable measure of cell viability; here long-term viability of cells was assessed using clonogenic cell survival assay. Under these exposure conditions, approximately one half of reversibly permeabilised cells lost their proliferative ability. This suggests that reversing permeabilisation is insufficient to ensure that cells maintain their proliferative ability. Cell viability depends not only on the recovery of the cell membrane following ultrasound treatment but also on other factors: ultrasound can induce cell apoptosis, increase intracellular ion concentration (Ando et al. 2006), and damage proteins on cell membranes (Kawai and Iino 2003), all of which may lead to cell death. The examination of different endpoints in in vitro studies may be important for optimization of sonoporation-mediated therapies. Although some applications may benefit from killing cells, most drug and gene delivery applications seek to maximize intracellular uptake while maintaining high levels of cell viability.

3.4.2 Microbubble agent Microbubbles significantly increase cell membrane permeabilisation; minimal intracellular delivery was observed in cells treated with ultrasound alone. Sonoporation is believed to be mediated through cavitation-related activities including stable oscillation and bubble disruption (Marmottant and Hilgenfeldt 2003, Ohl et al. 2006); therefore, it is not surprising that the addition

Chapter Three 54 of microbubbles can increase its efficiency. This finding is in agreement with other studies (Pislaru et al. 2003, Zarnitsyn and Prausnitz 2004) but also suggests that sonoporation efficiency depends on microbubble characteristics. The two microbubble-agents in this study contain similar perfluorocarbon gas cores. However, Definity has a lipid shell and a mean size range of 1.1-3.3 µm, whereas Optison has an albumin shell and a mean size range of 2-4.5 µm, with a lower number concentration. Microbubbles exposed at their resonant frequency undergo larger oscillations than at off resonance and the shear stress induced by microstreaming of oscillating bubbles influences membrane permeability (van Wamel et al. 2006). Microbubble disruption can result in shell rupture and release of gas, fragmentation and diffusion, all of which depend on microbubble characteristics and ultrasound exposure parameters (Bevan et al. 2008, Chen et al. 2002). In addition, the native bubble concentration of these agents is different: Definity has a 20- fold higher concentration than Optison. Several studies have indicated that the presence and type of microbubbles may be of crucial importance in sonoporation-mediated therapeutic applications, and that perfluorocarbon gas-filled microbubbles seem more efficient at inducing sonoporation than air-based ones (Blomley 2003, Li et al. 2003, Ward et al. 1999). It was also suggested that bubble-cell spacing is an important factor in predicting and controlling sonoporation (Miller et al. 2008, Ohl et al. 2006). Microbubble collapse close to the cell surface can cause severe damage and induce cell death (Miller et al. 2008), whereas, at longer distances, effects associated with inertial caviation may induce reversible sonoporation (Li et al. 2003). Future work should investigate the effect of microbubble size and bubble-cell spacing on sonoporation.

3.4.3 Limitations of the study This study has shown that not all reversibly permeabilised cells remain clonogenically viable. However, adherent cell lines such as KHT-C must form attachment bonds with the substratum, a process which involves components of the plasma membrane, in order to proliferate and form a cell colony. Furthermore, the optimum microbubble conditions were determined based on two agents that have been developed for clinical diagnostic use. Microbubble agents that are clinically approved and commercially available may be more practical for clinical translation of sonoporation. However, other agents, such as targeted microbubbles loaded with the therapeutic molecule, may be better suited for sonoporation-mediated therapies. Furthermore, the clinical relevance of the optimal microbubble concentrations, determined in this study, to in vivo conditions must be carefully evaluated. More experiments are required to determine optimal exposure conditions for in vivo applications.

Chapter Three 55 3.5 Conclusions We conclude that cell-impermeable macromolecules ranging in size from 10 kDa to 2 MDa can be delivered to the intracellular space of KHT-C cells at comparable efficiency. Microbubble type and concentration as well as acoustic pressure influence permeability efficiency. Approximately 30-50% of the resealed cells retained the capacity to proliferate. This study demonstrated that large macromolecules, up to 2 MDa in size, can be delivered with high efficiency to cells undergoing reversible permeabilisation, while maintaining the long-term viability in approximately half of the cells. Sonoporation mediated delivery of most therapeutic molecules may therefore not be limited by their size and by cell viability. Additional work is needed to identify comparable conditions in vivo.

Chapter Three 56 CHAPTER FOUR

Mechanism of cell sonoporation: Generation of transient sub- micron disruptions on the plasma membrane by ultrasound and microbubbles

4.0 Abstract The mechanism by which ultrasound and microbubbles increase the permeability of cell membranes is investigated through microscopic observations of the plasma membrane and uptake of cell-impermeable molecules. KHT-C cells in suspension were exposed to ultrasound pulses of 500 kHz centre frequency, 570 kPa peak negative pressure, 32 μs pulse duration, 3 kHz pulse repetition frequency and 5 s insonation time with and without Definity microbubble (3.3% v/v). Cell permeabilisation was assessed with 70 kDa FITC-dextran. Plasma membrane morphology was observed using electron microscopy. Cell permeability of 71% was achieved by ultrasound and microbubbles with the FITC-dextran added 60 s before ultrasound exposure. Permeability similar to control (0.5%) was achieved with ultrasound alone (2%), and with FITC-dextran added 60 s following termination of ultrasound and microbubble treatment (0.6%). Untreated cells exhibited continuous plasma membrane morphology as assessed with transmission electron microscopy (TEM). TEM images of cells treated with ultrasound and microbubbles revealed disruptions on the plasma membrane generally ranging from 30 nm to 100 nm and as large as 400 nm, immediately (i.e. within 2 s) following termination of ultrasound. The average disruption observed on the plasma membrane was 102±93 nm. No disruptions were observed on cells one minute after termination of ultrasound and microbubble treatment, and on cells treated with ultrasound alone. The presence of disruptions on the cell membrane was consistent with the intracellular uptake of FITC-dextran molecules. This study demonstrated that the biological mechanism underpinning sonoporation is the generation of transient sub-micron disruptions on cell membranes by ultrasound-activated microbubbles. (This chapter was submitted to Journal of Controlled Release)

Chapter Four 57 4.1 Introduction In this study, the mechanism by which ultrasound and microbubbles increase the permeability of cell membranes is investigated through microscopic observations of the plasma membrane and uptake of cell-impermeable molecules. The hypothesis guiding this study is that the mechanism of sonoporation is the generation of transient sub-micron disruptions on plasma membranes mediated by ultrasound and microbubbles. The objectives of this study are to directly observe bioeffects on the surface of cell membranes induced by ultrasound and microbubbles with electron microscopy methods and to correlate them with flow cytometry data on cell permeability.

4.2 Methods

4.2.1 In vitro cell model KHT-C cells in suspension were used (Section 2.2.1).

4.2.2 Ultrasound exposure system The ultrasound exposure apparatus (Section 2.2.2) was used to expose cells to optimal ultrasound conditions: 500 kHz pulse centre frequency, 570 kPa peak negative pressure, 32 μs pulse duration, 3 kHz pulse repetition frequency and 5 s insonation time, based on maximizing cell permeability (Chapter Two).

4.2.3 Ultrasound microbubble agent Definity agent was used in this study (Section 2.2.3). A volume of 50 μL ofDefinity microbubbles, corresponding to a volume concentration of 3.3% v/v, was added to the cell suspension 30 s before exposing the sample to ultrasound.

4.2.4 Reversible permeabilisation and PI-viability: PR and VPI Following ultrasound exposure, cell permeability and viability were assessed using 70 kDa FITC-dextran molecules and flow cytometry. Non-viable cells were identified by the presence of propidium iodide within them (Section 2.2.5). Permeabilised cells were identified by the presence of 70 kDa FITC-dextran molecules within them (Section 2.2.5). The FITC-dextran was added either 60 s before or 60 s after exposing the cells to ultrasound. Each measurement of cell permeability and viability was repeated four times. The results are reported as the mean and standard error of the mean.

Chapter Four 58 4.2.5 Electron microscopy SEM (scanning electron microscopy, Hitachi S-3400N, Hitachi High Technologies America Inc, Pleasanton, CA, USA) images of cells were acquired at 1 000-to-40 000 times magnification. A volume of 1.2 mL glutaraldehyde fixative (5%) was added to 1.2 mL volume of cells in suspension immediately (that is, within 2 s), following termination of the ultrasound exposure. In preparing the samples for SEM, the cells remained in the fixative solution for five minutes, and were then centrifuged onto polylysine-treated slides using a Cytospin centrifuge (Shandon Cytospin 3 Cytocentrifuge, Thermo Scientific, MA) at 1000 rpm for two minutes. The cells were then dehydrated through an ethanol series washing process beginning with 25% and changing to solutions of 50%, 75% and 100%. The samples were sputter coated with gold- palladium prior to imaging with SEM. TEM (transmission electron microscopy (TEM, Hitachi H-7000, Hitachi High Technologies America Inc, Pleasanton, CA, USA) images of cells were acquired at 5 000-to-100 000 times magnification. The glutaraldehyde fixative (5%) was added to the cell suspension immediately, 60 s, and one and 24 hours following termination of the ultrasound exposure. In preparing the samples for TEM, the cells remained in the fixative solution for 15 minutes, and were then centrifuged at 2000 rpm for one minute to form a pellet. The fixative was removed and replaced with 100% ethanol. The samples were further stained with osmium tetroxide and prepared for TEM imaging.

4.3 Results 4.3.1 Intracellular uptake of FITC-dextran Permeability and viability of cells exposed to four different conditions is shown in Figure

4.1. Minimal effects were observed in cells treated with ultrasound alone. PR with ultrasound alone (2%) was similar to the no exposure (0.5%) condition. The addition of microbubbles significantly enhanced the ability of ultrasound to permeabilise cell membranes. A PR of 71% was achieved in combination with microbubbles; a similar permeability was achieved in a previous study (Karshafian et al. 2009). In these samples, the FITC-dextran marker was added before exposing the cells to ultrasound. However, when FITC-dextran molecules were added one minute following ultrasound exposure with microbubbles, no FITC-dextran stained cells were detected with flow cytometry. Cell permeability in these samples was comparable to that of untreated cells. Permeable cells became impermeable within one minute following termination of ultrasound and microbubble exposure. Cell viability was reduced by ultrasound and microbubbles (VPI=80%), whereas viability with ultrasound alone was similar to control (VPI=98%).

Chapter Four 59 Figure 4.1: Uptake of FITC-dextran during sonoporation and one minute following

termination of the treatment. Cell permeability (PR), the percentage of cells stained

with FITC-dextran and unstained with PI, and cell viability (VPI), the percentage of cells unstained with PI, are shown for four conditions: Untreated, ultrasound alone, ultrasound and microbubbles with FITC-dextran added 60 s before treatment, and ultrasound and microbubbles but with FITC-dextran added 60 s after terminating the ultrasound treatment. Intracellular delivery of FITC-dextran is achieved only

when present during treatment with ultrasound and microbubbles. VPI was reduced to 80% by ultrasound and microbubbles. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v.

Chapter Four 60 4.3.2 Disruption of plasma membrane Electron micrographs of KHT-C cells, shown in Figures 4.2 and 4.3, demonstrated two noticeable changes related to plasma membrane morphology. Firstly, the absence of ruffles on the external surface of the plasma membrane in cells exposed to ultrasound and microbubbles compared to untreated ones is apparent on the SEM images (Figure 4.2). Untreated cells have numerous ruffles (Figure 4.2a, 5000 times magnification), which are a normal morphological characteristic. SEM micrographs of cells treated with ultrasound and microbubbles (Figure 4.2c, 10 000 times magnification) suggest that the ruffles were removed from the plasma membrane. The cells were fixed immediately following ultrasound and microbubble exposure.

Figure 4.2: Pore-like structures and ruffles on scanning electron microscopy (SEM) images of cells. SEM image of untreated cell shows numerous ruffles (a), which are absent on ultrasound and microbubble treated cells (c); acquired at 5 000x and 10 000x magnifications. Pore-like structures are observed on untreated cells (~100 nm) (b) and on treated cells (~300 nm) (d); acquired at 40 000 x magnification. The arrow indicates a structure that resembles a pore. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v.

Chapter Four 61 Secondly, structures resembling pores on the plasma membranes of cells treated with ultrasound and microbubbles are visible (Figure 4.2d), (n=9). Pore-like structures up to 400 nm were observed on SEM images of cells acquired at 40 000 times magnification (Figure 4.2d). Pore-like structures were detected on SEM images of untreated cells as well, although, they were less than 150 nm in size (Figure 4.2b, 40 000 times magnification). It was inconclusive as to whether these pore-like structures detected with SEM were actual pores through which molecules could cross the plasma membrane and enter the cell. High-resolution TEM images of cells treated with ultrasound and microbubbles revealed localized sub-micron disruptions in the plasma membrane compared to untreated ones (Figure 4.3). TEM images of control cells demonstrated normal morphologic characteristics such as a continuous plasma membrane (Figure 4.3a). TEM images of cells exposed to ultrasound alone and fixed immediately did not show any breaks in plasma membrane continuity (Figure 4.3b). They were similar to control cells. TEM images of cells, which were fixed immediately following treatment with ultrasound and microbubbles, showed disruption of the plasma membrane (Figure 4.3c); the images lacked the dark line suggestive of membrane morphology. The size of the disruptions generally ranged from 30 nm to 100 nm, with few of the disruptions as large as 400 nm; a histogram of disruption size from 12 cells is shown in Figure 4.4. The average disruption size was 102±94 nm (mean ± standard-deviation). The TEM images of ultrasound and microbubble treated cells that were fixed one minute following termination of ultrasound did not contain disruptions on the plasma membrane (Figure 4.3d). The cells seemed to have repaired the disruptions to their plasma membranes within a minute. In addition, the membranes of cells fixed at one hour and 24 hours following ultrasound and microbubble exposure resembled that of control cells. Furthermore, the contrast of the TEM images in ultrasound and microbubble treated cells was lower compared to control cells.

Chapter Four 62 Figure 4.3: High-resolution transmission electron microscopy (TEM) images at 80 000 times magnification of (a) control cells, (b) cells treated with ultrasound alone and fixed immediately, (c) cells treated with ultrasound and microbubbles and fixed immediately; and (d) cells treated with ultrasound and microbubbles and fixed one minute following ultrasound exposure. Disruptions on the plasma membrane are observed in (c) which are absent in (a), (b) and (d). Black arrow indicates a disruption on the plasma membrane. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v.

Chapter Four 63 Figure 4.4: Histogram of disruption size generated on the membrane of cells by ultrasound and microbubbles. Disruptions in the range of 30-to-100 nm were observed with few of the disruptions as large as 400 nm. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=570kPa, f=500kHz, T=5s, and Definity =3.3% v/v. Based on analysis of12 TEM images of cells treated with ultrasound and microbubbles and fixed immediately following termination of the treatment. No disruptions were observed on the TEM images of cells before treatment and one minute following termination of the treatment.

Chapter Four 64 4.4 Discussion

4.4.1 Mechanism of sonoporation This study demonstrated that transient sub-micron disruptions on the plasma membrane are responsible for the increased cell permeability by ultrasound and microbubbles. Microscopic observations showed the disruptions and their transient nature. Disruptions, generally within 30 to 100 nm, were generated on the plasma membrane of cells treated with ultrasound and microbubbles. The disruptions were short lived; they recovered within a minute following termination of ultrasound. The electron microscopy observations were consistent with the flow cytometry measurements. Intracellular delivery of FITC-dextran molecules was achieved when the permeability fluorescent marker was available in the extracellular space during the treatment with ultrasound and microbubbles. Cells recovered or at least re-established the barrier function of the membrane within a minute after terminating the treatment. The plasma membrane became impermeable to FITC-dextran molecules within a minute. The concept that sonoporation mediates delivery of cell-impermeable molecules through membrane disruptions has been suggested in other studies, however, with disruptions of micron scale (Mehier-Humbert et al. 2005, Schlicher et al. 2006). In this study, sub-micron scale disruptions were observed with electron microscopy. Studies on uptake of macromolecules support this finding. Molecules up to 56 nm could be delivered at high efficiency (~70% permeabilisation) (Karshafian et al. 2009), whereas delivery of larger molecules appeared to be less efficient (Mehier-Humbert et al. 2005). This study suggests that sub-micron size disruptions are most likely responsible for sonoporation mediated delivery of cell-impermeable molecules to live cells. The mode of uptake is considered to be through passive diffusion (Liang et al. 2004). A favourable concentration gradient across the plasma membrane causes molecules present in the extracellular space to diffuse into the cytoplasm through the disruptions formed on cell membranes. Ideally, subsequent studies should generate 3D TEM images of cells treated with ultrasound and microbubbles and measure size and spatial distribution of the disruptions on the plasma membrane. Electron microscopy provided snapshots of the plasma membrane disruption immediately and one minute after termination of ultrasound. It was found that cells repaired disruptions on the plasma membrane within a minute. This finding is consistent with the concept that sonoporation is a transient phenomenon. Cells usually repair their membrane within minutes, however, it was suggested that the recovery rate of the cell membrane may be dependent on the ultrasound exposure conditions (Deng et al. 2004, Kawai and Iino 2003, Mehier-Humbert et al. 2005, Okada et al. 2005,

Chapter Four 65 Pan et al. 2005). Real-time measurements of trans-membrane voltage demonstrated an increase in trans-membrane current across the plasma membrane during ultrasound and microbubble exposure, and recovery to a pre-treatment level following termination of ultrasound treatment (Deng et al. 2004). Cells exposed to ultrasound at low duty cycles repaired their membrane faster, based on faster recovery of trans-membrane current (Pan et al. 2005). The process through which pores on the plasma membrane are resealed may involve exocytosis of vesicles that fuse with each other and with the actin cytoskeleton to form repair patches where the plasma membrane has been disrupted (Schlicher et al. 2006). This process requires energy and Ca2+ ions (Deng et al. 2004). Other bioeffects of ultrasound and microbubble exposure on cells, such as removal of ruffles from the external surface of membranes and modification of cell surface smoothness were observed (Duvshani-Eshet et al. 2006, Mehier-Humbert et al. 2005).

4.4.2 Disruption induced by ultrasound-activated microbubbles Microbubbles were essential in inducing sonoporation, under the exposure conditions of this study. Morphological bioeffects on cell membranes including disruptions and removal of ruffles were induced by ultrasound and microbubbles. Ultrasound alone was not sufficient to disrupt the plasma membrane and induce uptake of FITC-dextran molecules. These cells appeared undisturbed, comprising a continuous plasma membrane structure similar to untreated cells. Cell permeabilisation was achieved through effects associated with stable microbubble oscillation such as microstreaming (Marmottant and Hilgenfeldt 2003, van Wamel et al. 2006), and with microbubble disruption such as the generation of shock waves and microjets (Ohl et al. 2006). A clinically relevant characteristic of sonoporation is that it can induce transient disruptions of biological membranes at will, potentially in a controlled manner, to facilitate intracellular delivery of cell-impermeable molecules. The molecular size of therapeutic agents may not be a limiting factor in sonoporation mediated therapies, however, the transient nature of sonoporation necessitates that molecules of the drug be in close vicinity to target cells during treatment to be effectively delivered. This may be accomplished by active targeting of microbubbles loaded with drugs to surface molecules (Lindner and Kaul 2001) or by pushing the microbubbles closer to the endothelial cell surface with ultrasound pulses (Ferrara 2008), and consequently releasing the therapeutic agent with ultrasound pulses.

Chapter Four 66 4.4.3 Limitations of the study In this study, the mechanism of sonoporation was investigated based on exposing in vitro cells in suspension to ultrasound and microbubbles, and correlating flow cytometry analysis with electron microscopy observations. Membrane disruption was considered to be the underpinning mechanism of sonoporation which allowed cell-impermeable molecules to enter the cell. However, such disruptions may involve components of the cytoskeleton system that were not investigated here. Furthermore, in vivo, cells are integrated together and with the extracellular matrix to form a tissue. Mechanical forces applied to a cell are distributed not only over the plasma membrane and the cytoskeleton but also over the extracellular matrix. The biological repair mechanism in vivo may most likely involve other aspects not present in single cells. Elucidation of the sonoporation mechanism in vivo and involvement of the cytoskeleton require further investigation.

4.5 Conclusions The mechanism underpinning sonoporation of mammalian cells was found to be the generation of transient sub-micron disruptions on the plasma membrane induced by ultrasound and microbubbles. Untreated cells exhibited normal continuous plasma membrane morphology. Ultrasound and microbubbles generated disruptions generally ranging from 30 to 100 nm on the plasma membrane of cells, allowing intracellular delivery of cell-impermeable molecules. Cells re-established the barrier function of the membrane and the disruptions were patched within a minute following termination of the ultrasound treatment. The dynamics of membrane disruption and subsequent repair observed with electron microscopy was in agreement with flow cytometry measurements of the ability of cell-impermeable molecules to cross the plasma membrane. Sonoporation was induced by ultrasound-activated microbubbles. No disruptions on the plasma membrane were observed and no permeabilised cells were detected in ultrasound alone treated cells. This study may be applied to guide the development of sonoporation-mediated therapies.

Chapter Four 67 CHAPTER FIVE

Dependence of sonoporation on cell cycle: Enhanced effect during later stages - Work in progress

5.0 Abstract In this study, we investigated the relationship between sonoporation and cell cycle phase. KHT-C cells in suspension were exposed to ultrasound at 500 kHz pulse centre frequency, 32 μs pulse duration, 3 kHz pulse repetition frequency and two minute insonation time at different peak negative pressures (125 and 570 kPa) in the presence of Definity microbubbles (3.3% v/v).

Following ultrasound treatment, PR and VPI were measured in the major phases of the cell cycle

(G1, S and G2/M) using fluorescent markers (PI, FITC-dextran and Hoechst 33342) and flow cytometry. Sonoporation of cells depended on cell cycle phase. Cells in the later stages were more prone to being permeabilised but even more to being killed by ultrasound and microbubbles.

Higher therapeutic ratio, TRP, of 5.3 was achieved in G1 compared to S and G2 phases. This may be related to differential biomechanical properties of cells such as size and viscoelasticity that change with cell cycle. This is the first study, to the best of our knowledge, which demonstrated a relationship between sonoporation and cell cycle phase.

Chapter Five 68 5.1 Introduction The dependence of sonoporation on the cell cycle phase is investigated here. The interaction of ultrasound-activated microbubbles with the plasma membrane generated disruptions through which cell permeability was enhanced allowing macromolecules to enter cells. The plasma membrane appears to be the primary site of action in sonoporation, hence its characteristics may be important in sonoporation. In the first three studies (Chapters Two to Four), the cell suspension model contained cells in different phases of the cell cycle. A cell, during its life cycle, replicates its DNA and undergoes cell division. This process can be subdivided into four steps: G1, S, G2 and

M. During G1 phase, a cell is prepared to replicate its DNA, which occurs during S phase. During

G2 phase, a cell is prepared for cell division, which occurs during M (mitosis) phase (Fahnestock et al. 1989). We hypothesize that the efficiency of sonoporation depends on the cell cycle phase. The objectives of this study are to measure cell permeability and viability in the major phases of the cell cycle (G1, S and G2/M) and to quantify the amount of molecules delivered to the intracellular space of cells at different ultrasound pressures.

5.2 Methods

5.2.1 In vitro cell model KHT-C cells in suspension were used in this study (Section 2.2.1).

5.2.2 Ultrasound exposure system The ultrasound exposure apparatus (Section 2.2.2) was used to expose cells to 500 kHz pulse centre frequency, 32 μs pulse duration, 3 kHz pulse repetition frequency and two minutes insonation time at two peak negative pressures (125 and 570 kPa).

5.2.3 Ultrasound microbubble agent Definity agent at 3.3% v/v was used during ultrasound treatment (Section 2.2.3).

5.2.4 Reversible permeability and PI-viability: Cell cycle phase

Cell permeability (PR) and viability (VPI) were measured in the major phases of the cell cycle (G1, S and G2/M) using three fluorescent markers – PI, FITC-dextran and Hoechst 33342 (Invitrogen Inc, Carlsbad, CA) – and LSR II flow cytometry (BD LSR II, BD Biosciences

Chapter Five 69 International, San Jose, CA). Prior to ultrasound exposure, FITC-dextran molecules (70 kDa) were added to the cell suspension (Section 2.2.5). Following ultrasound exposure, the cell suspension was rinsed and resuspended in 600 μL volume of PBS. Next, the anti-fluorescent dye (A-889) was added to the cell suspension; the samples were placed at 37 °C for 15 minutes. Thereafter, a 10 μL volume of Hoechst 33342 at a concentration of 1 mg/mL was added to the samples and incubated at room temperature for 30 minutes. Prior to flow cytometry analysis PI was added to the cell suspension (Section 2.2.5) to act as a marker of the non-viable cells. FITC-dextran marker assessed cell permeability induced at the time of the sonoporation exposure. Hoechst 33342 marker, a cell-permeable molecule that stains the DNA of viable cells, identified the cell cycle phase; the fluorescent intensity histogram was analysed using FlowJo software. The mean fluorescent intensity of FITC-dextran (IF) was determined in the whole population and each cell cycle phase. A total of 30,000 cells were counted per sample. Each measurement was repeated four times. The results are reported as the mean and standard error of the mean.

5.2.5 Measures of cell sensitivity: DPI and PE Cell sensitivity, in a specific cell cycle phase, to ultrasound and microbubble exposure was assessed using two indices: DPI and PE. DPI is defined as the number of cells killed by the treatment relative to untreated control:

VVCNTLP− I DPI = Equation 5.1 VCNTL

where VCNTL is the mean PI-viability of untreated control, and VPI is the PI-viability of treated samples. DPI was used as a measure of cell sensitivity to being killed by ultrasound and microbubbles with respect to cell cycle phase.

PE is defined as the ratio of permeabilised to PI-viable cells:

PR PE = Equation 5.2 VPI

where PR is the number of reversibly permeabilised cells and VPI is the number of PI-viable cells.

PE was used as a measure of cell sensitivity to being permeabilised by ultrasound and microbubbles.

Chapter Five 70 5.2.6 Therapeutic Ratio: TRP

The therapeutic ratio, TRP, was determined in the whole and each cell cycle phase using Equation 5.3.

PR Equation 5.3 TRP = VVCNTLP− I

where VCNTL is the mean PI-viability of untreated control, VPI is PI-viability and PR is cell permeability.

5.3 Results Sonoporation of KHT-C cells depended on cell cycle phase. Cells in later stages were more sensitive to the bioeffects induced by ultrasound and microbubbles. However, cells in the later stages were more prone to being killed than permeabilised resulting in higher TRP levels in the early stages of the cell cycle.

5.3.1 Cell PI-viability

The effect of cell cycle phase on PI-viability is shown in Figures 5.1 and 5.2. VPI decreased with acoustic pressure, as expected, but a larger fraction of cells in later stages of the cell cycle were killed with ultrasound and microbubbles (Figure 5.1). The number of cells killed varied statistically with the cell cycle phase; more cells were killed in G2 compared to G1 phase (Figure 5.2). A non-parametric statistical test, Mann-Whitney Rank Sum in MATLAB (The MathWorks, MA), was used to assess significant differences in the number of cells killed in each phase of the cell cycle, with p values less than 0.05. This effect was independent of acoustic pressure, under the exposure conditions of this study.

5.3.2 Reversible permeability The effect of cell cycle on cell permeability is shown in Figures 5.3 and 5.4. Cell permeability depended on acoustic pressure. PR’s similar to those reported in Chapter Two were achieved. However, a larger fraction of the cells in the later stages of the cell cycle were permeabilised (Figure 5.4). The differences observed in the number of permeabilised cells in each cell cycle phase were statistically significant with p values less than 0.05 (non-parametric statistical test, Mann-Whitney Rank Sum in MATLAB). Cells in later stages were more prone to being permeabilised by ultrasound and microbubbles.

Chapter Five 71 Figure 5.1: PI-viability (VPI) in the whole population, and G1, S and G2/M phases. VPI decreased with cell cycle stage and acoustic pressure.

40 G2

35 S G2 30 log 25 S G1 20 log DPI (%) 15 G1 10

0 125 570 Peak Negative Pressure (kPa)

Figure 5.2: The median and the 25th and 75th percentiles of PI-dead cells (DPI) in different phases of the cell cycle. The number of cells stained with PI molecules were

normalized with respect to the number of cells in controls in the whole population, G1,

S and G2/M phases. Statistically significant differences was measured in the number of cells killed with the phase of the cell cycle. More cells are killed in the later stages

(G2/M) compared to early stages (G1).

Chapter Five 72 Figure 5.3: Reversible cell permeability (PR) measured at varying acoustic pressures

(0, 125 and 570 kPa) in the whole population, and G1, S and G2/M phases. The number of cells stained with 70 kDa FITC-dextran and unstained with PI molecules, indicating

reversible permeabilisation (PR) are shown in the major phases of the cell cycle.

1 G2 S log 0.9 G1

0.8

0.7 G2 0.6 S PE Ratio 0.5 log G1 0.4

0.3

0.2 125 570 Peak Negative Pressure (kPa) Figure 5.4: The median and the 25th and 75th percentiles of the number of permeabilised

cells - efficiency ratio (PE) - in different phases of the cell cycle. PE, the ratio of the number of reversibly permeabilised cells with respect to the number of viable cells

remaining following treatment in whole population, G1, S and G2/M phases. A statistically

significant higher EP levels are achieved at the later stages.

Chapter Five 73 5.3.3 Therapeutic Ratio: TRP The therapeutic ratio in different phases of the cell cycle is shown in Figure 5.5. Cells in the early stages of the cell cycle achieved higher TRP levels. Although a larger fraction of the cells in the later stages of the cell cycle were permeabilised but even more were killed, resulting in lower TRP levels in later stages. A maximum TRP of 5.3 was achieved in G1 phase at 570 kPa.

A comparable TRP level of ~2 was achieved in S and G2 phases independent of acoustic pressure amplitude.

Figure 5.5: Therapeutic ratios, TRP , the number of permeabilised cells divided by the

number that are killed by the same exposure, of the whole population, G1, S and G2/M

phases are shown. Lower TRP , values were achieved in the later stages of the cell cycle.

Chapter Five 74 5.3.4 Intracellular uptake Fluorescence intensity of FITC-dextran in permeabilised cells with respect to cell cycle phase is shown in Figure 5.6. IF depended on cell cycle phase. Higher IF was achieved in later cell cycle stages; IF of 168 was achieved in G2 compared to 98 in G1 phase. This effect was less pronounced at lower pressures; IF of 63 was achieved in G1 compared to 82 in G2 at 125 kPa. The fluorescence intensity levels were not corrected for the cell size.

Figure 5.6: Fluorescence intensity of FITC-dextran delivered to the intracellular space

of cells in the whole population, G1, S and G2/M phases are shown. Higher intensities were achieved at the later stages of the cell cycle and higher acoustic pressures.

Chapter Five 75 5.4 Discussion This study demonstrated that sonoporation depends on cell cycle phase. Cells in later stages were more prone to being permeabilised but even more to being killed by ultrasound and microbubbles (Figures 5.2 and 5.4). This is the first study, to the best of our knowledge, which demonstrated a relationship between sonoporation and cell cycle phase. A similar dependence was observed in electroporation studies (Brunner 2002, Delgado-Cañedo et al. 2006, Golzio et al. 2002). Cells most affected by electroporation were those in advanced cell cycle phase. The dependence of sonoporation efficiency on cell cycle stage may be important inin vivo studies.

5.4.1 Relationship between sonoporation and cell cycle Cell size appears to be a factor in the dependence of sonoporation on cell cycle. Cells double their volume during cell division, resulting in a ~26% increase in the radius of a cell based on a spherical cell model. Early investigations on the mechanism of ultrasound-induced cell death showed that the largest cells in the population were especially sensitive to being killed by ultrasound (Thacker 1973). Furthermore, cell lysis induced by ultrasound was greater in larger cells (Miller and Battaglia 2003). However, cells undergo changes in membrane biomechanical properties and cytoskeleton during the cell cycle. For example, microviscosity of cells was maximum during cell division (de Laat et al. 1977). The question of whether rigid membranes are more prone to biomechanical perturbations induced by ultrasound microbubbles remains to be determined. Future work should investigate the dependence of sonoporation on the biomechanical properties of cell membranes.

5.4.2 Limitations of the study This study has several limitations. An in vitro cell suspension model of a single cell type was used to show the dependence of sonoporation on cell cycle phase. The cell cycle phase was identified using fluorescent molecules that bind to the DNA. However, the ultrasound and microbubble exposure conditions used in this study were based on maximizing cell permeability of a heterogenous cell population where majority of the cells were in G1 phase (~70%). This situation suggests the importance of using homogeneous populations for sonoporation studies. Future work will continue to investigate the dependence of sonoporation on cell cycle phase in other cell types. As well, cells will be synchronized to specific cell cycle phases and then exposed to ultrasound.

Chapter Five 76 5.5 Conclusions Sonoporation efficiency of in vitro cells depends on cell cycle phase. Cells in different stages of the cell cycle demonstrated varied sensitivity to ultrasound and microbubble exposure. Cells in the later stages were more prone to being permeabilised but even more to being killed by ultrasound and microbubbles. This may be related to differential biomechanical properties of cells such as size and viscoelasticity that change with cell cycle.

Chapter Five 77 CHAPTER SIX Summary and Future Directions

6.1 Summary and Conclusions

The goal of the work described in this thesis was to improve the understanding of the sonoporation process. Cells in suspension were exposed to ultrasound and microbubbles – a total of 97 exposure conditions. Following the treatment, the uptake of cell-impermeable molecules over a range of 10 kDa to 2 MDa, the long-term viability of cells following uptake, and the microscopic structure of plasma membranes were assessed using flow cytometry, colony assay and electron microscopy.

Based on the studies described in this thesis, we concluded the following (Figure 6.1): • Sonoporation is induced by ultrasound and microbubbles under the exposure conditions of the studies. • Sonoporation efficiency can be controlled through ultrasound exposure parameters in combination with microbubble characteristics. The exposure conditions can be tailored to achieve high sonoporation efficiency or to maintain high PI-viability but achieve lower sonoporation efficiency: Cell permeability of ~70% with ~25% cell death versus permeability of ~35% with ~2% cell death. • Microbubble disruption is a necessary but insufficient indicator of ultrasound-induced permeabilisation. Cell permeability and viability did not correlate with bubble disruption. Cell permeability as high as 70% and as low as 2% were achieved under ultrasound exposure conditions that disrupted ~99% of the bubbles. Minimal intracellular delivery was observed in cells treated with ultrasound alone (~2%).

Chapter Six 78 • Sonoporation efficiency may not be predicted by a single exposure parameter. Cell viability and membrane permeability depended non-linearly on the ultrasound and the microbubble conditions. Specifically, we demonstrated that acoustic energy may not serve as unifying parameter to predicting sonoporation efficiency. • Molecular size of common therapeutic drugs are unlikely to be a limiting factor in sonoporation-mediated therapies. Molecules up to 2 MDa in size can be delivered at high efficiency while maintaining cell viability. • Sonoporated cells can remain clonogenically viable. Following uptake of cell-impermeable molecules by ultrasound and microbubbles, approximately half of the cells were able to proliferate. Long-term viability depended on ultrasound and microbubble exposure conditions. • The mechanism of sonoporation is the generation of transient sub-micron disruptions on the plasma membrane by ultrasound and microbubbles allowing molecules at least up to 2 MDa to enter the intracellular space. Generally, disruptions between 30 and 100 nm were observed on plasma membranes. These disruptions were short lived; they recovered within one minute. • Sonoporation efficiency depends on the cell cycle phase. Cells in later stages of the cell cycle were more prone to being permeabilised and even more to being killed by ultrasound and microbubble exposure.

Chapter Six 79 Conclusion: Sonoporation Studies

Cell bubble

Before ultrasound and microbubble exposure

Cell-impermeable molecule Start ultrasound exposure

Membrane was disrupted (~100 nm) & Intracellular delivery of 10-to-2000 kDa was achieved in 70-80% of cells

Stop ultrasound exposure

Membrane resealed within one minute & ~50% of permeabilised cells remained clonogenically viable

Figure 6.1: Conclusions of this thesis: 1) Mechanism of sonoporation is the generation of transient sub-micron disruptions on cell membrane, 2) Sonoporation can be controlled by exposure conditions, 3) Cell-impermeable molecules can be taken up at high efficiency, 4) microbubble disruption is a necessary but insufficient indicator of ultrasound-induced permeabilisation, and 5) Long-term viability of sonoporated cells is possible.

Chapter Six 80 6.2 Future Directions The studies presented in this dissertation improved our understanding of the sonoporation process, however, additional work is still required. Future extension of this study can be in different directions. A parameter space of ultrasound and microbubble mediated permeabilisation of cell membranes was established based on experiments conducted in a single cell type (KHT-C cell). Sonoporation experiments on different cell lines including endothelial cell lines would be beneficial in determining whether the conclusions of the thesis may be generalized. Preliminary experiments were done using two endothelial cell lines (Section 6.2.1). The results reported in this thesis demonstrate that in in vitro cell models sonoporation efficiency can be tailored through ultrasound and microbubble exposure parameters. Future sonoporation experiments should also assess the effect of ultrasound and microbubble exposure in in vivo models. Preliminary experiments were done on in vivo blood vessels of fertilized chicken eggs (Section 6.2.2). Under the exposure conditions of the studies described in this thesis, permeabilisation of the cell membrane was induced by ultrasound-activated microbubbles that resulted in the disappearance of the majority of microbubbles within 0.8-to-20 μm range. However, the acoustic behavour of microbubbles during sonoporation remains to be determined. Microbubbles undergoing substantial oscillation may be determined through presence of sub-harmonic and ultra-harmonic signals (passively detected). Bubble disruption through inertial cavitation may be identified through passive detection of broadband acoustic signals. In addition, the microbubbles responsible for inducing sonoporation should be identified and associated with their acoustic response. This investigation may help explain the relationship between sonoporation efficiency and ultrasound exposure conditions. Furthermore, in the studies described in this thesis, the microbubbles and the cells were assumed to be uniformly distributed in the suspension model. Future studies should investigate the effect of the distance between a bubble and a cell on sonoporation efficiency. Furthermore, the fate and acoustic response of small microbubbles (< 0.5 μm) in inducing sonoporation should be investigated. The mechanism underpinning sonoporation-mediated uptake is the disruption of the plasma membrane barrier. However, whether the disruptions generated are non-specific or associated with specific structures of the cell membrane remains to be determined. In addition, the cytoskeleton, which reinforces the plasma membrane, is most likely affected during sonoporation. Future experiments will investigate the relationship between the viscoelastic properties of cells membranes and sonoporation efficiency. However, in addition to the transient membrane disruption mechanism

Chapter Six 81 for ultrasound induced delivery, biological internalization processes such as endocytosis has been shown to facilitate intracellular delivery at low ultrasound exposure conditions (Juffermans et al. 2008, Meijering et al. 2009). Furthermore, reversing permeabilisation, which was assessed through the exclusion of PI molecules, is insufficient to ensure that cells remain clonogenically viable. Understanding the process through which cells die may aid in improving efficiency of viable sonoporation. Preliminary results of ultrasound and microbubbles on in vitro endothelial cells and in vivo blood vessels of chorioallantoic membrane (CAM) are reported. The selection of the exposure conditions was guided by the studies described in this thesis.

6.2.1 Sonoporation of in vitro endothelial cells The effect of ultrasound and microbubble exposure on endothelial cells in suspension was investigated using the ultrasound exposure apparatus and flow cytometry methods specified in Chapter Two (2.2 Methods). Two endothelial cell lines were used: HUVEC (Human Umbilical Vein Endothelial Cells, a primary cell line) and EA.h96 (transformed endothelial cell line). The cells were maintained as monolayers in tissue culture flasks, harvested and suspended at a final concentration of 1.2x106 cells/mL and a volume of 1.5 mL. Prior to ultrasound exposure, a volume of 20 μL FITC-dextran (70 kDa) and Definity microbubbles (3.3% v/v) were added to the cell suspension. The endothelial cells were exposed to 500 kHz pulse centre frequency, 32 μs pulse duration, 3 kHz pulse repetition frequency and two minutes insonation time at two peak negative pressure amplitudes (125 and 246 kPa). Following ultrasound exposure, PR and VPI were determined with BD FACSCalibur flow cytometry. The TRR levels were calculated for the two cell lines. Each measurement was repeated four times and the results were reported as the mean and standard error of the mean. Sonoporation of endothelial cells was achieved with ultrasound and Definity microbubbles.

Cell permeability and viability is shown in Figure 6.2. PR increased and VPI decreased with peak negative pressure. Maximum PR’s of 63±1% and 58±2% were achieved at VPI’s of 76±1% and 61±1% with HUVEC and EA.h96 cells, respectively. The therapeutic ratio is shown in Figure

6.3. A maximum TRR of 2.6±0.2 was achieved in HUVEC cells compared to 1.5±0.1 in EA.h96 cells. Preliminary results suggest that sonoporation of HUVEC endothelial cells can be achieved at efficiency levels similar to KHT-C cells (Chapter Two).

Chapter Six 82 Figure 6.2: The effect of peak negative pressure on cell viability of two endothelial cell lines. The percentages of viable cells are shown with respect to peak negative pressure: Cell viability decreases with peak negative pressure. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-250kPa, f=0.5MHz, insonation time=2 minutes, and Definity 3.3% v/v.

Figure 6.3: Therapeutic ratio is shown for varying peak negative pressure. The therapeutic ratio is optimum at 246kPa peak negative pressure. Exposure conditions: PD=32μs, PRF=3kHz, Pneg=0-250kPa, f=0.5MHz, insonation time=2 minutes, and Definity 3.3% v/v.

Chapter Six 83 6.2.2 Sonoporation of in vivo CAM blood vessels The effect of ultrasound and microbubble exposure on blood vessels was investigated in vivo using the chorioallantoic membrane (CAM) of fertilized chicken eggs (Strain B, Shaver Poultry Breeding Farms, Cambridge, ON, Canada) between days 14 and 17. Prior to ultrasound exposure, the outer hard shell was removed to create a small rectangular window on the side and a larger one at the air pocket (Figure 6.4a). Paraffin oil was used to improve optical visualization and prevent drying of the CAM. A volume of 50 μL FITC-dextran (10 kDa) was injected into the circulation. Within a few minutes and following a bolus injection of 25 μL volume of Definity microbubbles, the CAM was exposed to ultrasound bursts every 2 s for two minutes. The ultrasound burst consisted of 32 μs pulse duration, 3 kHz pulse repetition frequency, 570 kPa peak negative pressure and 500 kHz pulse centre frequency for 50 ms. A schematic diagram and a photograph of the ultrasound exposure apparatus are shown in Figure 6.4b-c. It consisted of a customized water tank with 8

(a) (b) Microscope Transducer

Acoustic Signal Amplifier Window AWG Signal Generator Egg Acoustic Path (c) Figure 6.4: (a) The chorioallantoic membrane (CAM) of a fertilized chicken egg at day 15 with the shell removed. Blood vessels are optically visible. (b) A photograph depicting the ultrasound setup with the microscope unit on top. (c) A schematic diagram of the ultrasound exposure system.

Chapter Six 84 mm diameter acoustic window, a transducer (500 kHz) with its focus aligned with the acoustic window, and the electronic systems including waveform generator and power amplifier specified in Chapter Two. The CAM pocket was filled with saline, placed directly underneath the acoustic window and exposed to ultrasound. During the ultrasound exposure, the CAM was optically observed in real time through the acoustic window with a dissecting microscope (Wild M3B), and images were captured with a digital camera (Nikon DS 70) every few seconds. Following the exposure, the CAM layer was excised and the leakage of FITC-dextran molecules was assessed with a fluorescent microscope (Leica DM IL); fluorescent images of treated and untreated CAM were acquired. CAM blood vessels were significantly affected by the application of low frequency (500 kHz) ultrasound at 570 kPa peak negative pressure in the presence of Definity microbubbles. Optical images of the CAM acquired with the dissecting microscope before, during and after the ultrasound-and-microbubble exposure are shown in Figure 6.5. Preliminary observations indicated that large microvessels (100 to 300 μm) were vasoconstricted within 30-60 s, whereas smaller microvessels collapsed. Light and fluorescent microscopy images of the dissected CAM pre- and post-ultrasound acquired with an inverted microscope are shown in Figure 6.6. Small microvessels lost their integrity and collapsed. Larger microvessels appeared to be deformed but remained intact, just as was observed with the dissecting microscope. Fluorescent microscopy images of the dissected CAM showed that the fluorescent molecules leaked from blood vessels in ultrasound and microbubble treated areas of the CAM. Preliminary results indicate that sonoporation of blood vessels can be achieved with ultrasound and microbubbles. The in vivo CAM model combined with a real time fluorescent microscopy system may allow investigation of the sonoporation process in intact blood vessels and real time.

Chapter Six 85 Control: Prior to ultrasound expsoure Ultrasound Treated: Post exposure

(a) (c)

During ultrasound exposure: 30 s During ultrasound exposure: 5 min

(b) (d) 1 mm

Figure 6.5: (a) Control photograph of CAM vasculature perfused with microbubbles and fluorescent marker prior to ultrasound treatment. (b) Perfused vasculature 30 seconds into treatment. (c) CAM vasculature after five minutes of treatment; noticeable effect of vessel constriction is observed. (d) A photograph of CAM after treatment. Vessel constriction and disappearance of capillaries is apparent compared with control photograph.

Chapter Six 86 Control: Prior to ultrasound expsoure Ultrasound Treated: Post exposure

(a) (c)

Control: Fluorescent Image Ultrasound Treated: Fluorescent Image

(b) (d) 500 μm

Figure 6.6: Fluorescence microscopy images of excised CAM layer vasculature. (a) Control vessels from area untreated by ultrasound. (b) Fluorescence image of control vessels. (c) Ultrasound treated area. (d) Fluorescence image of ultrasound treated area.

Chapter Six 87 6.3 Final Comments The clinical utilization of sonoporation in drug and gene delivery applications still requires an improved understanding of sonoporation in in vivo models. Specifically, understanding of the underlying mechanism of sonoporation at the membrane level and the acoustic behaviour of microbubbles responsible for permeabilising but maintaining clonogenic viability must be improved. The development of imaging and therapeutic ultrasound integrated systems will further aid the development of sonoporation-mediated applications. The successful development of sonoporation-mediated drug and gene therapeutic applications in humans will most likely be a major advance in treating diseases including cancer.

Chapter Six 88 References

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References 98 APPENDIX Awards and Publications during doctoral period: (* related to thesis).

A. Awards 1. Award Of Excellence (Silver) in Cancer Research Category at CIHR National Health Research Poster Competition, 2007.* 2. Participant in the CSHRF (Canadian Student Health Research Forum) 2007: Nominated by the University of Toronto as being “within the top 2% of doctoral students in the health sciences” to participate in CIHR National Health Research Poster Competition, Winnipeg, June 2007.* 3. Merit Prize, International Society for Therapeutic Ultrasound (ISTU), Oxford, UK, 2006.* 4. Young Investigator Award (Bronze Prize), The 11th Congress of the World Federation on Ultrasound and Medicine (WFUMB), Seoul, Korea, 2006.* 5. CIHR Doctoral Award (2005-2008): Canadian Institute of Health Research, Canada.*

B. Publications

Peer-reviewed Journal Papers 1. Doria AS, Karshafian R, Moineddin R, Mohanta A, et al. Contrast-enhanced triggered harmonic sonography for assessment of periarticular hemodynamic changes in experimental arthritis. Pediatric Radiology, vol 36 (12), 1242-51, 2006. 2. Hirokawa T, Karshafian R, Pavlin CJ, Burns PN. Insonation of the eye in the presence of microbubbles: Preliminary study of duration and degree of vascular bioeffects – work in progress. Journal of Ultrasound in Medicine, Vol 26 (6), 731-38, 2007.* 3. Bevan PD, Karshafian R, Burns PN. Quantitative measurement of ultrasound disruption of polymer-shelled microbubbles. Ultrasound Med. Biol, vol 33 (1), 1777-86, July 2007.* 4. Cherin E, Brown J, Masoy SE, Shariff H, Karshafian R, Williams R, Burns PN, Foster FS. Radial modulation imaging of microbubble contrast agents at high frequency. Ultrasound Med. Biol, vol 34 (6), 949-62, 2008. 5. Bevan PD, Karshafian R, Burns PN. The influence of fragmentation on the acoustic response from shrinking bubbles. Ultrasound Med Biol, vol 34 (7), 1152-62, 2008.* 6. Needles A, Goertz DE, Karshafian R, Cherin E, Brown AS, Burns PN, Foster FS. High- frequency subharmonic pulsed-wave Doppler and color flow imaging of microbubble contrast agents. Ultrasound Med Biol, vol 34 (7), 1139-51, 2008. 7. Karshafian R, Bevan PD, Williams R, Samac S, Burns PN. The Effect of Acoustic Exposure Parameters on Cells and Bubbles: Cell Membrane Permeabilisation and Microbubble Disruption. Ultrasound in Medicine in Biology, Vol. 35 (5), 847-860, 2009.*

Awards and Publications 99 8. Karshafian R, Sanya S, Bevan PD, Burns PN. Microbubble mediated sonoporation of cells: Clonogenic viability and influence of molecular size on uptake. Ultrasonics (submitted). * 9. Karshafian R, Burns PN. Mechanism of cell sonoporation: Generation of transient sub-micron disruptions on the plasma membrane by ultrasound and microbubbles. J of Controlled Release (submitted).* 10. Hudson J, Karshafian R, Burns PN. Quantification of flow using ultrasound and microbubbles: A disruption replenishment model based on physical principles. Ultrasound in Medicine & Biology (conditionally accepted).

Conference Proceedings 11. Karshafian R, Samac S, Banerjee M, Bevan PD, Burns PN. Ultrasound-induced uptake of different size markers in mammalian cells. IEEE UFFC Proc 2005: 13-16.* 12. Needles A., Brown AS, Karshafian R, Burns PN, Foster FS, Goertz DE. High frequency subharmonic pulsed-wave doppler and color flow imaging of microbubble contrast agents. IEEE UFFC Proc 2005:629 – 632. 13. Karshafian R, Bevan PD, Samac S, Burns PN. The effect of acoustic exposure parameters on cell membrane permeabilisation by ultrasound and microbubbles. International Society of Therapeutic Ultrasound (ISTU), Oxford, UK, August 2006.* 14. Hudson MJ, Karshafian R, Burns PN. Quantification of Flow using Ultrasound and Microbubbles: A Disruption-Replenishment Model Based on Physical Principles. Proc IEEE UFFC Ultrasonics Symposium, 2006. 15. Shariff H, Bevan P, Kharshafian R and Burns P. Radial Modulation Imaging: Raising the frequency for contrast imaging. Proc. IEEE, Ultrasonics Symposium, Vancouver, Oct 2006. 16. Karshafian R, Bevan PD, Samac S, Burns PN. The Effect of Acoustic Exposure Parameters on Cell Membrane Permeabilisation by Ultrasound and Microbubbles. 6th International Symposium on Therapeutic Ultrasound (AIP Conference Proceedings), vol 911, 498-504, 2007.* 17. Cherin E, Brown J, Shariff H, Karshafian H, Williams R, Burns PN, Foster FS, Masoy SE. Radial modulation imaging of microbubbles at high frequency. Proc IEEE Ultrasonics Symposium, 888-91, 2007. 18. Karshafian R, Samac S, Bevan PD, Czarnota GJ, Burns PN. The dependence of sonoporation on cell cycle phase: Enhanced effect during G2 and S-phase. Proc IEEE Ultrasonics Symposium, 2005-07, 2007.* 19. Spraque MR, Cherin E, Karshafian R, Goertz DE, Foster SF. Acoustic Characterisation of Individual Targeted Microbubbles with High-Frequency Ultrasound. Proc IEEE Ultrasonics Symposium, 2008. 20. Goertz DE, Karshafian R, Hynynen K. Antivascular Effects of Pulsed Low Intensity

Awards and Publications 100 Ultrasound and Microbubbles in Mouse Tumors. Proc IEEE Ultrasonics Symposium, 2008. 21. Goertz GE, Karshafian R, Kullervo Hynynen. Acoustic emissions associated with ultrasound stimulated microbubble modulation of tumor blood flow. Proceedings IEEE International Ultrasonics Symposium 2009 (submitted). 22. Karshafian R, Burns PN. Ultrasound and microbubble mediated generation of transient pores on cell membranes in vitro. Proceedings IEEE International Ultrasonics Symposium 2009 (submitted). 23. Karshafian R, Giles A, Burns PN, Czarnota GJ. Ultrasound-activated microbubbles as novel enhancers of radiotherapy in leukemia cells in vitro. Proceedings IEEE International Ultrasonics Symposium 2009 (submitted).

Abstracts 24. Karshafian R, Bevan PD, Burns PN. Microbubble Potentiated Changes in Cell Permeability and Viability. IEEE International Ultrasonics, Ferroelectrics, and Frequency Control Joint 50th Anniversary Conference. August 23-27, 2004.* 25. Bevan PD, Karshafian R, Matsumura M, Ticker G, Burns PN. An Acoustic Study of Disruption of Polymer Shelled Bubbles. IEEE International Ultrasonics, Ferroelectrics, and Frequency Control Joint 50th Anniversary Conference. August 23-27, 2004.* 26. Wang FM, Karshafian R, Burns PN. A Real-Time Clinical Ultrasound Contrast Dosimeter with Adaptive Algorithms. IEEE International Ultrasonics, Ferroelectrics, and Frequency Control Joint 50th Anniversary Conference. August 23-27, 2004. 27. Burns NP, de Jong N, Bevan DP, Boukaz A, Karshafian R, Ticker G. Looking at and Listening to Breaking Bubbles: A Correlative Optical/Acoustical Study of Experimental Polymer/Air Agents. RSNA 2004. 28. Karshafian R, Samac S, Banerjee M, Bevan PD, Burns PN. Ultrasound-induced uptake of different size markers in mammalian cells. IEEE International Ultrasonics Symposium. Rotterdam, The Netherlands, 2005.* 29. Karshafian R, Samac S, Banerjee M, Bevan PD, Burns PN. Ultrasound-induced uptake of different size markers in mammalian cells. IEEE Ultrasonics Symposium, 2005.* 30. Needles, A.; Brown, A.S.; Karshafian, R.; Burns, P.N.; Foster, F.S.; Goertz, D.E.; High frequency subharmonic pulsed-wave doppler and color flow imaging of microbubble contrast agents. IEEE Ultrasonics Symposium, 2005. 31. Shariff H, Bevan P, Kharshafian R and Burns P. Radial Modulation Imaging: Adual frequency ultrasound imaging technique for microbubble contrast. European Symposium on Ultrasound Contrast Imaging, p64-66, Rotterdam, Jan- 2006. 32. Hudson MJ, Karshafian R, Burns PN. Contrast Quantification of Flow: Compensating

Awards and Publications 101 for the non-uniform ultrasound beam profile. The Eleventh European Symposium on Ultrasound Contrast Imaging, January 2006 (Young Investigator Award). 33. Bevan PD, Karshafian R, Burns PN. Acoustic response of shrinking bubbles. The11th European Symposium on Ultrasound Contrast Imaging. Rotterdam, The Netherlands. January 2006.* 34. Hirokawa T, Karshafian R, Burns PN. Ultrasound-Induced Permeability in an in-vivo rabbit eye. AIUM 2006.* 35. Shariff H, Bevan P, Kharshafian R and Burns P. Imaging Tumour Blood Flow using Radial Modulation Imaging, A New Ultrasound Contrast Imaging Technique, Ontario Center of Excellence for Breast Cancer Imaging Research, INO Symposium, p162, Toronto, April-2006. 36. Shariff H, Bevan P, Kharshafian R and Burns P. Imaging Breast Cancers using Radial Modulation Imaging, a new ultrasound contrast imaging technique. Centre for Research in Women’s Heath, 5th Annual Symposium, p26-27, Toronto, May-2006. 37. Karshafian R, Bevan PD, Samac S, Banerjee M, Burns PN. Sonoporating Cells for Drug Delivery: Effect of Pulse Bandwidth and Duty-Cycle. The 11th Congress of the World Federation on Ultrasound and Medicine (WFUMB), Seoul, Korea, 2006.* 38. Shariff H, Bevan P, Kharshafian R and Burns P. Radial Modulation Imaging: Raising the frequency for contrast imaging. IEEE International Ultrasonics Symposium, Vancouver, Oct 2006 (IEEE Student Award). 39. Hudson MJ, Karshafian R, Burns PN. Quantification of Flow using Ultrasound and Microbubbles: A Disruption-Replenishment Model Based on Physical Principles. IEEE International Ultrasonics Symposium, 2006. 40. Karshafian R, Bevan PD, Samac S, Burns PN. The effect of acoustic exposure parameters on cell membrane permeabilisation by ultrasound and microbubbles. International Society of Therapeutic Ultrasound (ISTU), Oxford, UK, August 2006.* 41. Tang A, Kim TK, Heathcote J, Guindi M, Jang HJ, Burns PN, Karshafian R, Wilson SR. Hepatic vein transit times using a microbubble agent to predict severity of hepatic disease non-invasively. RSNS 2006. 42. Bevan PD, Karshafian R, Samac S, Burns PN. Acoustic characterization of Definity disruption and implications for drug delivery. The 12th European Symposium on Ultrasound Contrast Imaging. Rotterdam, The Netherlands. January 2007.* 43. Karshafian R, Czarnota GJ, Giles A, Burns PN. Ultrasound and microbubble potentiated sensitization of cells to ionizing radiation. AACR, California, USA. April 2007. 44. Karshafian R, Samac S, Williams R, Giles A, Bevan PD, Czarnota GJ, Burns PN. Cancer Therapy Enhanced by Ultrasound and Microbubbles. CSHRF (Canadian Student Health Research Forum), Winnipeg, Manitoba, Canada. June 6, 2007.*

Awards and Publications 102 45. Karshafian R, Giles A, Burns PN, Czarnota GJ. Ultrasound Microbubble Potentiated Sensitization of Cells and Tumours to Radiotherapy. Radiotherapy and Oncology, vol 84, S27, 2007. CARO-COMP, Toronto, Canada. 46. Karshafian R, Samac S, Bevan PD, Czarnota GJ, Burns PN. The dependence of sonoporation on cell cycle phase: Enhanced effect during G2 and S-phase. IEEE International Ultrasonics Symposium. October 2007.* 47. Cherin E, Brown J, Shariff H, Karshafian H, Williams R, Burns PN, Foster FS, Masoy SE. Radial modulation imaging of microbubbles at high frequency. IEEE International Ultrasonics Symposium. October 2007. 48. Thind A, Karshafian R, Teitelbaum A, Goertz DE, Abdulla P, Strauss B, Foster FS. Ultrasound contrast agents for accelerating collagenase activity in Chronic Total Occlusion. The 13th European Symposium on Ultrasound Contrast Imaging. Rotterdam, The Netherlands. January 2008. 49. Papanicolau N, Azrif M, Karshafian R, Giles A, Sadeghian A, Kolios MC, Czarnota GJ. Conventional-Frequency Ultrasound Detection of Apoptosis In Vivo. American Institute of Ultrasound in Medicine (AIUM) 2008. 50. Czarnota GJ, Karshafian R, Giles A, Banihashemi B, Lee J, Burns P. Microbubble and Ultrasound Enhancement of Radiation-Induced Tumor Cell Death In Vivo. American Institute of Ultrasound in Medicine (AIUM) 2008. 51. Czarnota GJ, Karshafian R, Giles A, Banihashemi B, Lee J, Burns P. Microbubble and Ultrasound Enhancement of Radiation-Induced Tumor Cell Death In Vivo. Annual Meeting of the American Association of Cancer Research, April 2008. 52. Papanicalau N, Lee J, Banihashemi B, Debeljevic B, Ranieri S, Azrif M, Karshafian R, Giles A, Kolios MC, Sadeghian A, Czarnota GJ. Novel Low-Frequency Ultrasound Detection of Apoptosis in vitro and in vivo. Ultrasound Imaging and Tissue Characterization Symposium (UITC) 2008. 53. Czarnota GJ, Lee J, Karshafian R, Banihashemi B, Chu W, Cho C, Kolios MC. High- frequency detection of cell death: Assessment of Chemotherapy, Radiotherapy, Photodynamic Therapy and Novel Microbubble-Therapy Effects. Ultrasound Imaging and Tissue Characterization Symposium (UITC) 2008. 54. Papanicolau N, Muhammad A, Karshafian A, Giles A, Sadeghian A, Kolios MC, Czarnota GJ. Conventional-frequency ultrasound detection of apoptosis in vivo. AIUM, March 2008. 55. Czarnota GJ, Karshafian R, Giles A, Banihashemi B, Justin L, Burns PN. Microbubble and ultrasound enhancement of radiation-induced tumour cell death in vivo. AIUM, March 2008. 56. Lee JW, Karshafian R, Banihashemi B, Caissie A, Giles A, Burns P, Czarnota G. Novel

Awards and Publications 103 enhancement of cancer responses to radiation utilizing ultrasound-activated microbubbles: Histopathological-treatment outcomes. International Journal of Radiation Oncology Biology Physics, vol 72 (1), S706, 2008. 57. Lee JW, Karshafian R, Banihashemi B, Caissie A, Czarnota G. Ultrasound microbubble- potentiated enhancement of tumour response to radiation: Preliminary results. Annual Research Day, University of Toronto, May 2008. 58. Thind AS, Karshafian R, Teitelbaum A, Goertz DE, Adbulla P, Strauss F, Foster S. Ultrasound contrast agents for accelerating collagenase activity in Chronic Total Occlusion. Heart and Stroke Joint Annual Research Day, May 2008. 59. Lee J, Karsafian R, Banihashemi B, Caisse A, Giles A, Giles A, Burns PN, Czarnota GJ. Enhancement of Prostate Cancer Responses to Radiation Utilizing Ultrasound-Activated Microbubbles. ASTRO 2008. 60. Goertz DE, Karshafian R, Kullervo H. Antivascular effects of pulsed low intensity ultrasound and microbubbles in murine tumors. 8th ISTU Conference, September 2008. 61. Thind AS, Karshafian R, Teitelbaum A, Goertz DE, Adbulla P, Strauss F, Foster S. Softening the matrix: The use of ultrasound mediated microbubbles to enhance collagenase activity in Chronic Total Occlusion. Imaging Network Ontario, September 2008. 62. Sprague MR, Cherin E, Karshafian R, Goertz DE, Foster FS. Probing the acoustic response of individual targeted microbubbles with high-frequency ultrasound. Imaging Network Ontario, September 2008. 63. Furukawa M, Hupple C, Nofiele J, Karshafian R, Cassie A, Giles A, Czarnota GJ. Doppler Ultrasound Imaging of Microbubble-Enhanced Radiation Responses in Tumours. Imaging Network Ontario, September 2008. 64. Sprague MR, Cherin E, Karshafian R, Goertz DE, Foster FS. Characterization of individual targeted microbubbles with high frequency ultrasound. World Molecular Imaging Congress, 2008. 65. Thind AS, Karshafian R, Teitelbaum A, Goertz DE, Adbulla P, Strauss F, Foster S. Ultrasound mediated microbubbles for accelerating collagenase activity in ex vivo model of Chronic Total Occlusion. Canadian Cardiovascular Society, October 2008. 66. Sprague MR, Cherin E, Karshafian R, Goertz DE, Foster FS. Acoustic characterization of individual targeted microbubbles with high-frequency ultrasound. IEEE Ultrasonics Symposium, November 2008. 67. Goertz DE, Karshafian R, Hynynen K. Modulating tumour blood flow with pulsed low intensity ultrasound and microbubbles. IEEE Ultrasonics Symposium, November 2008. 68. Czarnota GJ, Lee J, Giles A, Karshafian R, Banihashemi B. In vivo detection of tumor response to radiotherapy and new radiosensitization agent using quantitative non-invasive high frequency ultrasound. International Journal of Radiation Oncology Biology Physics,

Awards and Publications 104 Vol 72 (1), S683. 69. Cassie A, Al-Mahrouki A, Furukawa M, Karshafian R, Giles A, Lee J, Li Y, Wong S, Czarnota GJ. In vitro and in vivo vascular effects of novel radiosensitizing ultrasound- activated microbubbles. CARO (submitted).* 70. Goertz GE, Karshafian R, Kullervo Hynynen. Acoustic emissions associated with ultrasound stimulated microbubble modulation of tumor blood flow. IEEE International Ultrasonics Symposium 2009. 71. Karshafian R, Burns PN. Mechanism of Sonoporation: Ultrasound and microbubble mediated generation of transient pores on cell membranes in vitro. IEEE International Ultrasonics Symposium 2009. 72. Karshafian R, Giles A, Burns PN, Czarnota GJ. Ultrasound-activated microbubbles as novel enhancers of radiotherapy in leukemia cells in vitro. IEEE International Ultrasonics Symposium 2009. 73. Czarnota G, Karshafian R, Mahrouki AA, Giles A. Microbubble and Ultrasound Enhancement of Radiation-Induced Tumour Cell Death In Vivo. IEEE International Ultrasonics Symposium 2009. 74. Czarnota G, Karshafian R, Giles A, Mahrouki AA, Iradji Sara. Microbubble and Ultrasound Induction of Gene Expression Associated with Novel Radiation Enhancing Therapy. IEEE International Ultrasonics Symposium 2009. 75. Czarnota G, Karshafian R, Giles A, Al Mahrouki A. Ultrasound-activated microbubble enhancement of radiation response. ISTU (International Society of Therapeutic Ultrasound) 2009.

Awards and Publications 105 “The most exciting phrase to hear in science, the one that heralds the most discoveries, is not “Eureka!” but “That’s funny...” ”

Isaac Asimov

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