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Role of liposome mediated drug delivery and drug release in determining the therapeutic activity of liposomal… Lim, Howard J. 1999

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R O L E O F L I P O S O M E M E D I A T E D D R U G D E L I V E R Y A N D D R U G R E L E A S E IN D E T E R M I N I N G T H E T H E R A P E U T I C A C T I V I T Y O F L I P O S O M A L F O R M U L A T I O N S O F M I T O X A N T R O N E by H O W A R D J. LEVI B . S c , University of British Columbia, 1994 A T H E S I S S U B M I T T E D IN P A R T I A L F U L F I L L M E N T O F T H E R E Q U I R E M E N T F O R T H E D E G R E E O F D O C T O R O F P H I L O S O P H Y in T H E F A C U L T Y O F G R A D U A T E S T U D I E S D E P A R T M E N T O F P A T H O L O G Y A N D L A B O R A T O R Y M E D I C I N E We accept this thesis as conforming to the required standard T H E U N I V E R S I T Y O F B R I T I S H C O L U M B I A September, 1999 © Howard J. L i m , 1999 In presenting this thesis in partial fulfilment of the requirements for an advanced degree at the University of British Columbia, I agree that the Library shall make it freely available for reference and study. I further agree that permission for extensive copying of this thesis for scholarly purposes may be granted by the head of my department or by his or her representatives. It is understood that copying or publication of this thesis for financial gain shall not be allowed without my written permission. The University of British Columbia Vancouver, Canada DE-6 (2/88) A B S T R A C T Although liposomal accumulation at the target site is an important issue, the critical parameter defining the activity of a liposomal formulation is drug release, a factor that includes where, when, and how fast the therapeutic agent dissociates from the liposomal carrier. This point was investigated using two liposomal formulations of the anti-cancer drug mitoxantrone. Mitoxantrone was encapsulated via a pH gradient method in liposomes prepared o f 1,2 distearoyl-sn-glycero-3-phosphocholine (DSPC)/cholesterol (Choi) (55:45 mol ratio) or 1,2 dimyristoyl-sn-glycero-3-phosphocholine ( D M P C ) / C h o l (55:45 mol ratio), the latter exhibiting a greater rate o f drug release in vivo. Using a model of liver localized cancer consisting of B D F 1 mice inoculated with either P388 or L1210 cells intravenously (/.v.), it was demonstrated that a single dose of D M P C / C h o l mitoxantrone (10 mg/kg) administered i.v. resulted in 100% 60 day survival. In contrast, no long-term survivors were obtained in animals treated with free or D S P C / C h o l mitoxantrone. Drug levels in the liver were determined and demonstrate that greatest drug delivery was achieved with the D S P C / C h o l liposomal formulation. In an effort to address whether liposome mediated delivery or drug release is the dominant factor determining therapeutic activity, additional experiments examined the role of drug release at tumour sites where liposome accumulation is slow. As demonstrated in subcutaneous L S I 8 0 and A431 tumours grown on the backs of S C I D / R A G - 2 mice, the D M P C / C h o l formulation demonstrated greater activity in the L S I 8 0 tumour model and was as efficacious as the D S P C / C h o l formulation when treating A431 tumours. These data emphasize the importance of designing liposomal formulations that optimize drug biological availability rather than drug delivery. In an effort to understand factors that are important in governing the activity of D M P C / C h o l liposomal mitoxantrone used to treat liver localized disease, studies modulating liposomal accumulation in the liver were completed. Two methods were used to effect reductions in ii liposome delivery to the liver: the use of P E G modified lipids and hepatic mononuclear phagocyte system (MPS) blockade. Both methods reduced liposomal drug accumulation in the liver by a factor of 2 to 3 fold. A significant reduction in therapeutic activity was observed when PEG-modif ied lipids were incorporated into the D M P C / C h o l mitoxantrone formulation; however, M P S blockade did not affect anti-tumour activity. Long term survival (>60 days) was still observed in animals where hepatic M P S blockade effected elimination o f liver Kupffer cells. It is concluded that reductions in therapy observed for the PEG-modified D M P C / C h o l mitoxantrone are likely due to inhibition of cell binding and processing. Conversely it is suggested that the activity of the D M P C / C h o l mitoxantrone is dependent on cell processing, but the Kupffer cells do not play a significant role in this processing event. iii T A B L E O F C O N T E N T S ABSTRACT ii TABLE OF CONTENTS iv LIST OF FIGURES vii LIST OF TABLES x ABBREVIATIONS xi ACKNOWLEDGEMENTS xii DEDICATION xiv CHAPTER 1 INTRODUCTION 1 1.1 Foreword 1 1.2 Liposomes 4 1.2.1 Phospholipids 5 1.2.2 Cholesterol 8 1.2.3 Preparation of liposomes 9 1.2.3.1 Unilamellar vesicles (LUV and SUV) 10 1.2.4 Drug encapsulation 11 1.2.4.1 Passive entrapment (Figure 1.3) 13 1.2.4.2 Active entrapment (Figure 1.3, 1.4) 14 1.3 Liposomes as drug carriers 15 1.3.1 Liposomal anti-cancer drugs 18 1.3.2 Mitoxantrone 21 1.4 Biological fate of liposomes following intravenous administration 24 1.4.1 Barriers and compartments 24 1.4.2 Liposome serum protein interactions 26 1.4.2.1 Serum induced increases in membrane permeability 27 1.4.2.2 Serum protein binding and liposome elimination 28 1.4.3 Extravasation through blood vessels 31 1.4.3.1 Tumour vasculature 31 1.4.4 Role of the mononuclear phagocytic system (MPS) 33 1.4.4.1 Liposome accumulation in the liver 34 1.4.4.2 Decreasing liposome interactions with the MPS 37 1.4.5 Liposome extravasation 39 1.5 Dissociation of the active agent from the carrier: the critical parameter 40 1.5.1 Drug release - importance of drug type 42 1.5.2 Future considerations for the next generation of liposomes 44 1.5.3 Other methodology considerations 45 1.6 Thesis objectives and hypotheses 46 iv C H A P T E R 2 M A T E R I A L S A N D M E T H O D S 49 2.1 Materials 49 2.2 Preparation of liposomes 50 2.3 Transmembrane pH gradient loading of mitoxantrone 51 2.4 Transmembrane pH gradient loading of vincristine 52 2.5 Transmembrane p H gradient loading of doxorubicin 52 2.6 Preparation of E P C / C h o l clodronate liposomes 53 2.7 Microculture tetrazolium assay 53 2.8 In vitro characteristics of liposomal mitoxantrone 55 2.9 Plasma elimination and distribution studies 55 2.10 A431 and L S I 80 tumour accumulation and plasma elimination studies of liposomal mitoxantrone 56 2.11 Plasma elimination and biodistribution studies in M P S blockaded mice 57 2.12 Liposome mediated drug delivery to region o f tumour cell inoculation 57 2.13 Establishing the maximum tolerated drug dose and L1210 and P388 efficacy studies 58 2.14 Liposomal mitoxantrone anti-tumour efficacy using the human A431 and L S I 8 0 solid tumour models ' 59 2.15 Treatment of non-established L S I 80 and A431 tumours 60 2.16 Efficacy of liposomal mitoxantrone in the i.v. L1210 tumour model with and without M P S blockade 61 2.17 F4/80 staining of macrophages in the liver 61 2.18 Hepatocyte isolation 62 2.19 Confocal microscopy 62 2.20 Statistical analysis 63 C H A P T E R 3 I N F L U E N C E O F D R U G R E L E A S E C H A R A C T E R I S T I C S O N T H E T H E R A P E U T I C A C T I V I T Y O F L I P O S O M A L M I T O X A N T R O N E 64 3.1 Introduction: 64 3.2 Results 66 3.2.1 In vitro mitoxantrone uptake and release characteristics 66 3.2.2 In vivo plasma elimination of liposomal lipid and mitoxantrone 68 3.2.3 Acute toxicity of free and liposomal mitoxantrone 71 3.2.4 L1210 and P388 anti-tumour activity of free and liposomal mitoxantrone 74 3.2.4 Drug and liposomal lipid uptake in liver 77 3.3 Discussion 79 C H A P T E R 4 F A C T O R S A F F E C T I N G T H E T H E R A P E U T I C A C T I V I T Y O F L I P O S O M A L M I T O X A N T R O N E F O L L O W I N G I N T R A V E N O U S A D M I N I S T R A T I O N I N S C I D / R A G 2 M I C E B E A R I N G E S T A B L I S H E D H U M A N A431 A N D LSI 80 S O L I D T U M O U R S : D R U G R E L E A S E V E R S U S L I P O S O M E M E D I A T E D D R U G D E L I V E R Y 86 4.1 Introduction 86 4.2 Results 88 4.2.1 L i p i d and drug accumulation in solid L S I 8 0 and A431 tumours 88 4.2.2 Efficacy of single dose administration of liposomal and free mitoxantrone in established A431 and L S 180 human solid tumours 93 4.2.3 Efficacy of multiple dose administration of liposomal and free mitoxantrone in non-established A431 and L S I 80 human solid tumours 95 4.2.4 Drug accumulation in the region of tumour cell inoculation 99 4.3 Discussion 102 C H A P T E R 5 R O L E O F K U P F F E R C E L L S A N D L I P O S O M E M E D I A T E D D R U G D E L I V E R Y T O L I V E R I N G O V E R N I N G T H E E F F I C A C Y O F D M P C / C H O L L I P O S O M A L M I T O X A N T R O N E U S E D T O T R E A T L I V E R L O C A L I Z E D C A N C E R 109 5.1 Introduction 109 5.2 Results 112 5.2.1 Therapeutic activity of free and liposomal anti-cancer drugs given i.v. to mice bearing the L1210 /.v. tumour model 112 5.2.2 Reducing D M P C / C h o l liposomal mitoxantrone delivery to the liver 114 5.2.3 Influence of reducing liver mitoxantrone levels on the therapeutic activity of D M P C / C h o l mitoxantrone 115 5.2.4 Influence of hepatic M P S avoidance and elimination strategies on mitoxantrone release 119 5.2.5 Influence of hepatic M P S avoidance and elimination strategies on liposome distribution in the liver and on Kupffer cell depletion 121 5.3 Discussion 130 C H A P T E R 6 S U M M A R I Z I N G D I S C U S S I O N 136 6.1 Summary of results 136 6.2 Discussion 138 R E F E R E N C E S 144 vi LIST OF FIGURES Figure 1.1 Liposome target site accumulation 3 Figure 1.2 Structure of a phospholipid 6 Figure 1.3 Illustration of passive and active entrapment 12 Figure 1.4 Active drug entrapment in liposomes 16 Figure 1.5 Structure of mitoxantrone hydrochloride 22 Figure 1.6 Structure of a capillary: continuous endothelium type 25 Figure 1.7 Liposomal drug release 29 Figure 1.8 Diagram of a classic liver lobule 36 Figure 3.1 Effect of temperature on pH gradient loading of mitoxantrone into DSPC/Chol and DMPC/Chol liposomes 67 Figure 3.2 Release of mitoxantrone from DSPC/Chol (»)and DMPC/Chol (•) liposomes in HEPES buffered saline at 37 °C 69 Figure 3.3 Release of mitoxantrone from DMPC/Chol liposomes incubated with fetal bovine serum at 37°C for 24 hours 70 Figure 3.4 In vivo release of mitoxantrone from DSPC/Chol (•), DMPC/Chol (•) liposomes, and free mitoxantrone (A ) 72 Figure 3.5 Liver and spleen weights of untreated BDF1 mice (open bars) and BDF mice previously (7 days) injected i.v. with 104 L1210 cells (hatched bars) 75 Figure 3.6 Progression of the P388 i.v. tumour model in the liver 76 v i i Figure 3.7 Survival times o f B D F 1 mice injected with 104 L1210 cells (Panel A ) or 10 5 P388 cells (Panel B) i.v. via the lateral tail vein and treated with mitoxantrone 80 Figure 3.8 L i p i d and drug levels in the liver of mice after injection of D S P C / C h o l mitoxantrone ( • ) , D M P C / C h o l mitoxantrone(B), empty D M P C / C h o l liposomes ( • ) , and free mitoxantrone (A ) 81 Figure 4.1 Lip id and mtioxantrone accumulation in A431 and L S I 80 tumours in S C I D / R A G - 2 mice over a 96 hour time period 91 Figure 4.2 Percentage o f initial drug to lipid ratio of D S P C / C h o l mitoxantrone and D M P C / C h o l mitoxantrone after 48 hours in plasma 92 Figure 4.3 Efficacy o f D S P C / C h o l mitoxantrone, D M P C / C h o l mitoxantrone and free mitoxantrone in established L S I 8 0 and A431 solid tumours in S C I D / R A G - 2 M i c e 96 Figure 4.4 Drug accumulation at the site of tumour cell inoculation following i.v. administration o f free mitoxantrone or mitoxantrone encapsulated in D M P C / C h o l or D S P C / C h o l liposomes 101 Figure 5.1 Plasma elimination of D M P C / C h o l mitoxantrone liposomes and D M P C / C h o l / P E G Mitoxantrone liposomes 116 Figure 5.2 Drug accumulation in the liver versus therapeutic activity 117 Figure 5.3 Drug release of mitoxantrone from D M P C / C h o l liposomes and D M P C / C h o l / P E G liposomes. 120 Figure 5.4 F4/80 staining of Kupffer cells in the liver 123 Figure 5.5 Confocal imaging of biodistribution of D i l labeled D M P C / C h o l mitoxantrone liposomes and D M P C / C h o l / P E G mitoxantrone liposomes in the liver 125 Figure 5.6 Confocal imaging of biodistribution of D i l labeled D M P C / C h o l mitoxantrone liposomes with and without M P S blockade 1277 Figure 5.7 Drug delivery to hepatocytes 129 viii Figure 6.1 Future design of liposomes LIST OF TABLES Table 1.1. Transition temperature (T c ) of various combinations of acyl chain length, degree of saturation, and headgroup moiety 8 Table 1.2 Major antineoplastic agents evaluated in a liposomal drug carrier system 21 Table 3.1 L1210 anti-tumour activity of free and liposomal mitoxantrone in B D F 1 mice 78 Table 3.2 Area under the liposomal-lipid and mitoxantrone concentration-time curves obtained in tumours and plasma from 0 to 96 hours following i.v. administration of free and liposomal drug (10 mg/kg dose) in S C I D / R A G - 2 mice bearing established A431 and L S I 8 0 tumours 94 Table 4.1 Attributes o f the L S I 80 and A431 cell lines and their growth characteristics in S C I D / R A G - 2 mice 97 Table 4.2 Treatment of non-established A431 and L S I 80 subcutaneous human xenografts in S C I D / R A G - 2 mice 98 Table 4.3 Treatment of non-established L S I 80 subcutaneous human xenografts in S C I D / R A G - 2 mice... 100 Table 5.1 Therapeutic activity of free and liposomal formulations of doxorubicin, vincristine and mitoxantrone following a single i.v. injection in mice bearing the LI210 i.v. tumours 113 Table 5.2 IC 5o o f doxorubicin, vincristine, and mitoxantrone when incubated with L1210 cells for 24 hours. 3 114 Table 5.3 Influence o f PEG- l ip id incorporation and hepatic M P S blockade on the L1210 Anti-tumour Act iv i ty of D M P C / C h o l Mitoxantrone 119 x A B B R E V I A T I O N S % ILS percentage increase in lifespan. 3 [ H ] - C D E [ 3H] Cholesteryl hexadecyl ether A U C area under the curve B S A bovine serum albumin Choi cholesterol C R P C reactive protein Di l 1,1'-Dioctadecy 1-3,3,3',3'-tetramethy 1 indocarbocyanine perchlorate D M P C l,2-Dimyristoyl-sn-glycero-3-phosphocholine D O X doxorubicin D S P C l,2-Distearoyl-sn-glycero-3-phosphocholine D S P E distearoyl phosphatidylethanolamine E D T A ethylenediaminetetra-acetic acid E P C egg phosphatidylcholine F B S fetal bovine serum FITC fluorescein isothiocyanate g centrifugal force H B S H E P E S buffered saline H B S S Hank's balanced salt solution H E P E S A^-2-hydroxyethylpiperazine-A^-2-ethane-sulphonic acid I C 5 0 Concentration necessary for 50% inhibition of growth I.L.S. Increase in life span up. intraperitoneal i.v. intravenous L U V large unilamellar vesicles M A C membrance attack complex Mitox mitoxantrone M L V multilamellar vesicles mol mole M P S mononuclear phagocytic system M T T 3-(4, 5-dimethylthiazol-2yl)-2, 5-diphenyl tetrazolium bromide M T D maximum tolerated dose PBS phosphate buffered saline P E phosphatidylethanolamine P E G poly(ethylene glycol) P E G - P E poly(ethylene glycol)-modified phosphatidylethanolamine Q E L S quasielastic light scattering s.c. subcutaneous S.D. standard deviation S . E . M . standard error of the mean . S U V small unilamellar vesicles T c gel-liquid crystalline phase transition temperature Vine vincristine XI A C K N O W L E D G E M E N T S Is Life worth living? That depends on the liver. - Anonymous. Essentially it will be difficult to thank everyone who helped make it to this point since there have been so many people who supported me throughout this journey and I would like to say a heartfelt - Thank You! A very special thanks to Marcel for trusting me 1) with your house 2) with your horse 3) as a grad student! I've had a great time and learned more than I could possibly imagine. Thank you for 1) your tremendous support 2) convincing me to do a Ph.D. 3) teaching me the big picture to science and research and 2) showing me how to teach as well. (I hope I haven't corrupted too many kids on Bowen... .). And of course, thanks to Jlonka. I'm sure my thesis took up some of Marcel's time, which could have been better spent with you. Thanks to Lawrence for great laughs and for sharing your philosophy on science. Thanks also to Kirsten Skov, Tom Madden, and Paul Sorenson - What a great committee!!! Thanks for referring me to Marcel (Tom) and for the encouragement from my whole committee. Thanks to Gwyn Bebb for sharing his wisdom and kind words over some sort of dessert. O f course I can't forget all those past and present who have lived down in the deep pit known as the Lower main floor of the B . C . Cancer Research Centre. Thanks to Troy, Murray, Shane, and Sandy who I must have annoyed endlessly with questions those first few months. Many thanks to Dana, Natashia, Nancy, and Rebecca for their help with the animals and cell cultures. An d of course Kelly, Frances, Hafiza, Raj, Peirrot, Lynn, Spencer, Guoyang, Y . P . , Zaihui, Gary, Dody, Ellen, Dawn, the N H L , every restaurant on the Broadway strip, Safeway, the Q C group for their reagents and glassware, Don Phillips and Ernie Hildman of Analytical Cytology, Maryesse St. Louis for her work on the confocal microscope at U . B . C . and of course my poor summer students (who kept me honest): Julie, Daniel, and Anne. Many thanks to the Division of Medical Oncology for giving me opportunity to learn the clinical aspects of cancer as well as for allowing me access to the seminar room for all those committee meetings. O f course, I will also remember those who helped me survive the non-science parts of my degree: Lil l ian, Julie, and Penny. Special thanks to two ladies who made my life so much easier by washing my glassware: Maria and Shelley. Thanks to M r . and Mrs. Tan who also provided a home away from home and their words of encouragement. A nd of course all my friends (new and old) who put up with me all this time. Thanks to my Dad and to Auntie Maggie for all their support. P.S. Marcel - If you ever come up with liposomal A D H , I might be able to drink some beer. xii A special Thank You: I want to thank the teachers and professors who have made this whole journey exciting and most of all fostered my desire to learn. I would not be here i f it were not for the effort of these people Thank you very much. David Llyod George Elementary School M s . Mitchel l - Grade 2 Gilpin Elementary School M s . Nelson - Grade 4 M s . Home - Grade 5 and 6 M r . B e l l - Grade 7 Magee Secondary School: M s . M c C o l l - Drama and Act ing M r . Nakashima - Chemistry M r . Baptie and M r . Rollins - Biology M r . Jung, M r Cheeseman, and Mr . Mi l l e r - Mathematics and Calculus M r . Keenlyside - Social Studies University of British Columbia Undergraduate Studies \ Dr. Swanson - Calculus Dr. Orvig - Chemistry Dr. Goumeniouk - Pharmacology Dr. Norris - History of Medicine Dr. Blake - Biomechanics Graduate Studies Dr. M c M a n u s Dr. Al la rd Dr. Brooks Dr. Devine Dr. Gi lks Dr. Pritchard Dr. Skov Dr. Sorenson Dr. Madden Dr. Mayer Dr. Ba l ly xiii D E D I C A T I O N To my mom, Thank you for raising me to be the person I am today. Thank you for teaching me the value of learning both inside and outside the academic world. Your support, patience and love, has been truly remarkable. And To Lisa, Thank you for being there for those late night calls and panic attacks. Y o u are truly a great friend. Who knew I would need all those highlighters. Enjoy! xiv C H A P T E R 1 I N T R O D U C T I O N 1.1 Foreword In their present form, liposomal carriers primarily impact drug biological availability and this, in turn, results in a number of important therapeutic benefits. This includes the well-established reduction in toxicity for liposomal formulations of drugs such as the anti-cancer agents doxorubicin (Gabizon et al., 1982; Olson et al, 1982; Mayer et al., 1994;) and vincristine (Mayer et al., 1993; Boman et al., 1994; Kanter et al, 1994) and the anti-fungal agent amphotericin B (Graybill et al., 1982; Krause and Juliano, 1988). This reduced toxicity does not occur at the expense of therapeutic activity and, as a result, the therapeutic index of these drugs is improved through liposomal encapsulation. The reasons that liposomal drug carriers improve the therapeutic properties of an associated drug are not well understood. Pre-clinical studies suggest that free drug (drug released from liposomes) remains the biologically active agent and that therapeutic improvements arise from liposome mediated changes in drug circulation lifetime and tissue distribution (Hwang, 1987; Mayer et al, 1994; Gabizon and Martin, 1997). Reduced toxicity may be related to reduced availability of drug to sensitive tissues. For example, in the case of doxorubicin where cardiotoxicity represents a significant treatment limiting toxicity, reduced levels of drug in cardiac tissue are observed when the drug is administered in liposomal form (Gabizon et al, 1982; Olson et al, 1982). Furthermore, it is believed that therapy results from enhanced drug accumulation at the disease site and this is mediated by extravasation and localization of the drug loaded liposomal carrier ("passive targeting"). This appears to be a relatively general phenomenon in that liposomes preferentially accumulate in sites of inflammation (O'Sullivan et al., 1988), infection (Bakker-Woudenberg et al, 1992); and tumour growth (Richardson et al, 1979; Proffitt et al., 1983; Gabizon and Papahadjopoulos, 1988). Perhaps the most significant liposome characteristic to consider, in addition to the carrier's effect on drug delivery, is the carrier's drug retention attributes. This thesis is principally concerned with the importance of controlled drug release. Drug release attributes must, however, be considered in the context of when, where and how rapidly drug release occurs. The illustration shown in Figure 1.1 is useful in orienting the reader to the fundamental premise guiding the research described. It is believed that for an intravenously administered liposomal anti-cancer drug to be optimal it must possess different attributes depending on where the liposome is localized. While in the blood compartment the liposome should retain drug. This w i l l serve two purposes: 1) to minimize systemic exposure of free drug and 2) to maximize delivery of the liposomal drug to sites outside the blood compartment. The latter is typically a slow process and i f the drug release rates are too rapid, liposomes which have left the blood compartment may contain little drug. Once localized in the site of disease development, the liposomes ideally must undergo a transformation process resulting in drug release from the liposome. This chapter reviews how this model of liposome delivery was developed and the results presented in subsequent chapters support the contention that drug release combined with liposome-mediated changes in drug distribution work together to enhance the therapeutic activity of an associated anti-cancer drug. It is important to recognize that the research described in this thesis was developed using simple liposome formulations. In fact it is argued here that the primary advantage of using liposomal carriers, as opposed to other carrier technologies, is due to the fact that it is not complicated. Procedures for making physically and chemically well-defined liposomes as well as procedures for encapsulating certain drugs in liposomes such that extremely high trapping efficiencies and 2 Figure 1.1 Liposome target site accumulation To allow for passive accumulation to the target site appropriately designed liposomal carriers must be retained in the blood compartment for extended time periods (A). While in the blood compartment liposomes interact with the cells lining the endothelium (B) or with specific target cells (C). Passive targeting is dependent on the presence of altered vascular endothelium, alterations that permit extravasation of circulating macromolecules (D). Following extravasation liposomes can release drug while residing in the interstitial space (E) or can be taken up by tumour associated macrophages (F). The potential to achieve specific interactions with target cells (active targeting) through use of targeting ligands is also feasible (G). (Figure reproduced from Harasym, 1997) 3 high drug-to-liposome ratios have been developed and these procedures are reviewed in this introduction. Although many investigators are working towards improved technology, developing liposome carriers that exhibit modified surface features (Allen et al, 1989; 1991; Gabizon, 1992), targeting ligands (Leserman et al, 1981; Ahmad et al, 1993) and/or membrane fusing attributes (Holland et al, 1996b; Kirpoi t in et al, 1996), it is important to recognize that our understanding of the mechanisms governing the activity of simple liposome formulations is relatively poor. The research described in this thesis leads to a better understanding of how liposomes function as anti-cancer drug carriers and this information is essential for those interested in developing improved technology. 1.2 Liposomes Lipids can be extracted from natural sources (e.g. cellular membranes) and upon hydration orient into a spheres of bilayers, resulting in the formation of liposomes. First observed by Bangham et al. (1965), these multilamellar vesicles ( M L V ) consisted of concentric lipid bilayers separated by aqueous channels. The bilayer configuration arises due to the amphipathic nature of the l ipid; the hydrophilic head group and hydrophobic tail of the lipid molecule orient the lipid molecules such that the head groups face the aqueous environment and the fatty acyl chains are oriented toward one another. Liposomes were first used as model membrane systems because they form closed spheres which have a defined interior aqueous space separated by lipid bilayers, making them a valuable tool for study of the structural and functional role of lipids in the biological membrane. This included investigations of membrane fusion (Dunham et al., 1977; van Meer et al, 1985; Bailey and Cul l i s , 1994), membrane-protein interactions (Rogers and Strittmatter, 1975; Sogor and Z u l l , 4 1975; Bortoleto et al, 1998; Yamaji el al, 1998), complement activation (Devine el al, 1994), and multi-drug resistance (Shapiro and Ling , 1995). Liposomes are also used to study ion gradients and membrane permeability (Deamer and Nichols , 1983; Viero and Cul l is , 1990). This section w i l l focus on the two components, phospholipid and cholesterol, which typically are used in the preparation of liposomal drug delivery systems. 1.2.1 Phospholipids Phospholipids (or glycerophospholipids) consist of a glycerol backbone with a phosphate group esterified at the C(3) position and fatty acids esterified at the C ( l ) and C(2) positions as shown in Figure 1.2. Changes in the headgroup and/or in the fatty acids dictate the properties exhibited in the lipid bilayer. For example, liposomes containing the lipids phosphatidylserine, phosphatidylglycerol, phosphatidylinositol, and phosphatide acid w i l l have a negative surface charge at physiological pH. Liposomes with phosphatidylserine and phosphatidic acid are rapidly eliminated from the circulation following intravenous (i.v.) administration, in part because o f serum protein binding effects attributed to the negative charge (Moghimi and Patel, 1989). Since the ability of liposomes to move from the blood compartment to an extravascular site is dependent on maintaining a sufficient plasma concentration of liposomes for extended time periods, anionic phospholipids are not typically used when developing liposomal drug carriers. Instead there has been a focus on using zwitterionic phospholipids, in particular phosphatidylcholine. In addition to the importance of phospholipid head group charge, the acyl chain composition of the phospholipid can dramatically effect the characteristics of liposomes and their use as drug carriers. A key property of phospholipids is the temperature of the gel to liquid-crystalline phase 5 Figure 1 . 2 Structure of a phospholipid (Figure reproduced from Parr, 1995) o=p—o" A -C H j - C H — C H 2 — 0 A ' = A A==o A H 2 A H 2 I I CHo CHo 1 1 CH2 CHo 1 l C H 2 C H 2 A H 2 A H 2 A H 2 A H 2 A H 2 A H 2 A H 2 A H A H 2 C H Jl 1 CHo CHo 1 1 CHo CHo 1 1 CH2 CHj A H 2 A H 2 1 1 CHo CHo 1 1 CH3 C H 2 A H 2 - H e a d g r o u p G l y c e r o l b a c k b o n e A c y l c h a i n 'Neutral phospholipids C h o l i n e P h o s p h a t i d y l c h o l i n e ( P C ) « C H 2 C H 2 N + ( C H 3 ) 3 E t h a n o l a m i n e CUM+U P h o s p h a t l d y l e t h a n o l a m i n e ( P E ) ' - " 2 ^ " ^ ' "3 Negative phospholipids P h o s p h a t i d i c a c i d ( P A ) S e r i n e P h o s p h a t i d y l s e r i n e (PS) G l y c e r o l P h o s p h a t i d y l g l y c e r o l ( P G ) I n o s i t o l P h o s p h a t l d y l l n o s l t o l (PI) • « — C H 2 C H - N + H 3 Aoo" - — C H 2 C H ( O H ) C H 2 O H O H O H H O H / H O H H Saturated fatty acids L a u r i e ( 1 2 : 0 ) M y r i s t i c ( 1 4 : 0 ) P a l m i t i c ( 1 6 : 0 ) S t e a r i c ( 1 8 : 0 ) Unsaturated fatty acids P a l m i t o l e i c ( 1 6 : 1 A 9 ) O l e i c ( 1 8 : 1 A 9 ) L i n o l e l c ( 1 8 : 2 A 9 . 1 2 ) C H a f C H a J ^ H s C H C H z C H s C H J C H ^ T C O O H C H 3 ( C H 2 ) 1 0 C O O H C H 3 ( C H 2 ) 1 2 C O O H C H 3 ( C H 2 ) 1 4 C O O H C H 3 ( C H 2 ) 1 6 C O O H C H s f C H j J s C H ^ H f C H z j T C O O H C H s f C H ^ T C H ^ H f C H ^ T C O O H 6 transition (T c ) and this property is determined by both head group chemistry (Kruyff et al. 1973; Chowdry and Dalzie l , 1985) and acyl chain composition (McElhaney, 1982; Wang et al, 1997; Huang et al, 1993). The temperature where phospholipids undergo the transition from the gel to liquid-crystalline phase is referred to as the T c and the length and saturation o f acyl chains is a determining factor where the T c is observed. Typically, longer, more saturated acyl chains give rise to higher phase transition temperatures. The acyl chains are characterized by an order parameter "s" where s = 1 for no motion, and s = 0 for rapid isotropic motion. Below the T c , the acyl chains have a high "order" (s ~ 1), meaning that the chains are packed together in a frozen or "gel" phase where motion of the acyl chains is restricted. A t temperatures above the T c , the acyl chains are more fluid and less ordered in a "liquid-crystalline" phase. Longer acyl chains have increased order whereas unsaturated acyl chains disrupt packing and reduce the acyl chain order o f the membrane. The phospholipid head group can also affect the T c , as seen in Table 1.1. In general, membranes are more permeable to a variety of solvents and solutes at or above the T c than below (Bittman and Blau, 1972) and increased unsaturation or shorter acyl chains have been correlated with increased membrane permeability (Papahadjopoulos et al, 1973). A simple example of how the T c can be used in designing effective liposomal carriers concerns development o f what have been termed temperature sensitive liposomes (Weinstein et al. 1980; Magin et al, 1986). These liposomes are composed primarily of dipalmitolylphosphatidylcholine, which has a T c of 41°C. These liposomes can be induced to release entrapped contents by inducing local hyperthermia at regions where these liposomes accumulate following i.v. administration and result in increased drug availability. The studies described in this thesis also take advantage of differences in acyl chain composition to promote drug release. In particular, dimyristoyl- (T c = 24 °C) and distearoyl- (T c = 55 °C) phosphatidylcholine are used as the primary phospholipid components of the liposomes 7 Table 1.1 I Transit ion temperature (T c ) of various combinations of acyl chain length, degree of saturation, and headgroup moiety L i p i d Species Transit ion Temperature T c (°C) D i l a u r o y l P C ( 1 2 : 0 , 12:0) -1 Dimyristoyl P C (14:0, 14,0) 24 Dipalmitoyl P C (16:0, 16:0) 41 Distearoyl P C (18:0, 18:0) 55 Stearoyl, oleoyl P C (18:0, 18:1) 6 Stearoyl, linoleoyl P C (18:0, 18:2) -13 Dipalmitoyl P A (16:0, 16:0) 67 Dipalmitoyl PE(16:0 , 16:0) 63 Dipalmitoyl PS (16:0, 16:0) 55 Dipalmitoyl P G (16:0, 16:0) 41 characterized in this thesis. Differences observed in drug release from these liposomes are attributable, at least in part, to difference in permeability ascribed to acyl chain composition. 1.2.2 Cholesterol Cholesterol is the major neutral l ipid component of eukaryotic biological membranes and is composed of a rigid steroid ring and a polar 3-P-hydroxyl group. It orients itself with the hydroxyl group toward the lipid/water interface and the rigid steroid ring associated with the acyl chains. The flexible aliphatic tail extends into the membrane. Incorporation of cholesterol into the bilayer results in a decrease in the membrane order for phospholipids in the gel phase and increases the order of the membrane for lipids in the liquid-crystalline phase (De Kruyf f et al., 1973; Demel and de Kruyff, 1976). A t amounts above 7 mol %, the enthalpy of the gel to liquid crystalline phase transition is reduced until at 33 mol % and greater, the phase transition can no longer be detected (Hubbell and McConnel , 1971). Addition of cholesterol to unsaturated and saturated phosphatidylcholine (PC) membranes above their phase transition temperatures 8 decreases membrane permeability, while increasing the membrane permeability for membranes composed of saturated PC below the T g (Bittman and Blau, 1972). Cholesterol is an essential component of liposomes i f they are to be used as drug carriers. The presence of cholesterol at levels in excess of 30 mol% reduces serum protein binding (Patel et al, 1983; Semple et al, 1996). This, in turn, increases the circulation lifetime of the carrier (Kirby et al., 1980, Patel et al, 1983) and decreases release of entrapped contents (Fielding and Abra, 1992). The stabilizing role of cholesterol has been best illustrated by studies completed in Scherphofs laboratory (Scherphof et ai, 1978; 1979). These investigators demonstrated that liposomes prepared of dimyristoylphosphatidylcholine were completely "dissolved" when incubated with serum at the T c (24°C), an effect attributed to interactions with lipoproteins. Addit ion of cholesterol eliminated the serum-mediated destruction of these liposomes. 1.2.3 Preparation of liposomes Upon hydration of lipids, multilamellar vesicles ( M L V s ) are formed. These liposomes are heterogeneous and range in diameter from 1-10 microns. M L V s have proven to be of limited value for pharmaceutical applications, particularly those involving i.v. administration. These liposomes are rapidly eliminated from the plasma following injection due to their large size (Rahman et al, 1982). In addition, these liposomes tend to have a low trapped volume due to the tight packing o f the bilayers (Perkins et al, 1988). This trapped volume can increase with the incorporation of charged lipids that promote swelling of the liposomes due to the electrostatic replusion between the bilayers (Hope et al, 1986). In addition, methods that promote more efficient hydration of the lipids can also increase the trapped volume of M L V s . These methods would include reverse phase procedures (Szoka and Papahadjopoulos, 1978; 1980), dehydration-9 rehydration methods (Shew and Daemer, 1985), and those that use repeated freeze/thaw cycles (Mayer et al, 1985a; Ohsawa et al, 1985). Although the large size and heterogeneous nature of M L V s make them unsuitable for systemic applications, the steps used in the preparation of the M L V s define some of the attributes and the ease of manufacturing of the unilamellar liposomes that are commonly used for drug carrier applications. M L V precursors used in this thesis typically were subject to the freeze-thaw procedure to ensure equilibrium solute distribution and optimal trapped solute concentration. The latter term refers to circumstances where the trapped solute concentration is equivalent to the solute concentration used when hydrating the dried lipids (Mayer etal, 1985a). 1.2.3.1 Unilamellar vesicles ( L U V and S U V ) Small unilamellar vesicles (SUVs) range in size from 25-50 nm in diameter and are produced by sonicating M L V s or by forcing M L V s under high pressure through small openings. The latter refers to a method that originally used a French press (Barenholz et al, 1979) or as more recently developed for large-scale manufacturing of S U V s , an automatic high-pressure system called a Microfluidizer (Cheng et al, 1987). Although relatively easy to prepare and scalable to large (>10 L ) batch size, vesicles produced by these techniques tend to be unstable due to the curvature of their membranes and an associated propensity to fuse and form larger membrane structures. In addition, these systems tend to have small trapped volumes (<0.2 ul/umol) making them less suitable as drug carriers. Finally, following i.v. administration S U V s are small enough to penetrate the fenestrations that exist in the blood vessels of the liver (Hwang and Beaumier, 1986) and are accumulated in this organ at a faster rate then unilamellar liposomes that exhibit a mean diameter just 2- to 3- times larger. Combined, these properties make S U V s less useful as drug carriers. 10 Large unilamellar vesicles ( L U V ' s ) exhibit a mean diameter of between 50 and 400 nm and the majority of the vesicles consist of one bilayer enclosing an aqueous space. Many procedures have been described for the preparation of L U V s , but the most versatile and frequently utilized technique involves extruding M L V s through polycarbonate filters using high pressures of an inert gas (Olson et al, 1979; Hope et al, 1985; Mayer et al, 1986). This procedure forms a homogeneous population of unilamellar liposomes of well defined sizes depending on the pore size o f the filter used (50 nm -200 nm). L U V s are most suitable for drug delivery applications because o f their higher trapped volumes (1.5 to 10 ul/umole lipid), stability and pharmacokinetic/ biodistribution characteristics. This size has been found to be optimal for stability in the circulation as well as extravasation through vasculature. Unless otherwise indicated the remaining sections of this introduction refer to the preparation and in vitro/in vivo characterization of L U V s designed for intravenous applications as drug carriers. 1.2.4 Drug encapsulation There are essentially two techniques available for drug encapsulation: passive trapping and active trapping. These are illustrated in Figure 1.3. Passive trapping involves the addition of drug during the hydration of l ipid. The efficiency of this encapsulation procedure depends on the nature of the compound, where the level of hydrophobic compound association is governed by the capacity of the bilayer to incorporate the agent and the level of hydrophilic compound encapsulation is dependent on the aqueous trapped volume of the liposome used. Act ive trapping refers to techniques that involve addition of the therapeutic agent to pre-formed liposomes. 11 Figure 1.3 Illustration of passive and active entrapment Type Method Passive encapsulation of hydrophobic drug Drug added to solvents used to solubilize lipids or added during hydration of lipids Passive encapsulation of hydrophilic drug Drug added during hydration of lipids Active encapsulation of drugs exhibiting a protonizable amine function Generation of an ion gradient (e.g. pH gradient) followed by addition of selected drug 12 Compounds that are hydrophobic w i l l partition into the lipid bilayer. Alternatively techniques have been developed that rely on the chemical attributes of the drug and use of transmembrane ion gradients. 1.2.4.1 Passive entrapment (Figure 1.3) Passive entrapment of drugs is accomplished by the preparation of liposomes in the solution of the agent that is to be entrapped (Taylor et al, 1990). Use of this method generally results in poor drug entrapment and low drug-to-lipid ratios. For example, passive encapsulation of the anti-cancer drug doxorubicin results in a 4% trapping efficiency and a drug-to-lipid ratio of 0.004:1 (wt:wt) (Shinozawa et al, 1981). Trapping efficiency and drug-to-lipid ratio attributes are, of course, dependent on the aqueous trapped volume of the liposome as well as the lipid concentration when preparing the liposome. Mayer et al. (1986,), for example, demonstrated 80% trapping efficiency using liposomes extruded through 100 nm pore size polycarbonate filters. Considering the trapped volume of these liposomes is typically between 1.5 and 2.5 ul/umole lipid, 80% trapping efficiency can only be obtained by preparing the liposomes at high lipid concentration (up to 400 umol/ml). Hydrophobic drugs, such as cyclosporin A , are also entrapped in this manner (Ouyang et al, 1995). In this case, drug incorporation is dependent on the packing constraints of the drug in the membrane and the lipid characteristics. This procedure can result in high drug entrapment efficiency, but low drug-to-lipid ratios. Drugs of this class often exchange into other membranes rapidly and thus, in vivo the drug leaves the carrier quickly (Choice et al, 1995). 13 Liposomal anti-cancer drug formulations described in this thesis use transmembrane ion gradient based trapping methods (see next section), but it is important to recognize that these gradient techniques rely on passive trapping procedures to prepare the liposomes for drug loading. A s indicated above, the method should promote equilibrium solute distribution and the trapped volume of the liposome should be sufficient to insure an adequate trapping capacity (Boman et al, 1993). For these reasons, liposomes prepared for use in active drug loading procedures are typically generated by extrusion of frozen and thawed M L V s through 100 nm pore size polycarbonate filters. The lipid concentration used when preparing these liposomes is not as critical as that required for passive encapsulation of hydrophilic drugs. 1.2.4.2 Act ive entrapment (Figure 1.3, 1.4) The active trapping procedure is identified with any technique where drugs are loaded into preformed liposomes. For this reason any procedure using hydrophobic drugs that partition into the membranes of pre-formed liposomes can be defined as an active loading technique. It is more common, however, to associate active loading procedures with drugs that exhibit protonizable amine functions which can accumulate inside preformed liposomes exhibiting a transmembrane p H gradient (Mayer et al, 1985b; Madden et al, 1990; Mayer et al, 1993). The mechanism for accumulation (see Figure 1.4) is in response to a proton gradient where the interior of the liposomes have acidic p H . In the external environment, the neutral form of the weak base is membrane permeable and crosses the lipid bilayer. Once it enters the internal acidic environment, the weak base becomes protonated. The protonated form is then unable to permeate back across the lipid bilayer and is effectively "trapped" within the interior of the liposome. Assuming the p K a is the same on both sides of the membrane, the intravesicular and 14 external drug concentration can be derived from the Henderson-Hasselbach equation as: [ H A + ] i n / [ H A + ] o u t = [Fr] i n/[H +] 0 U 1 Therefore, a difference of 3 pH units between the exterior and interior of the liposome wi l l permit drug accumulation up to a maximum drug gradient of 10 3 fold higher inside versus outside. There are many advantages to the use of this procedure. First, this technique allows for trapping efficiencies approaching 100%. In addition, the rate of drug efflux is decreased by approximately 30-fold (Mayer et al, 1986) when compared to the same drug (doxorubicin) passively encapsulated in liposomes. Finally, provided the buffering capacity of the internal buffer has not been depleted, this trapping method works independently o f the starting drug to lipid ratio and can be used with almost any liposome formulation which is capable of maintaining an ion gradient. A s noted in Chapter 2, anti-cancer drug loaded liposomes where prepared using the p H gradient based loading procedure, where 300 m M citrate buffer (pH 4.0) was trapped inside. Many variations of the ion gradient based loading procedures have been developed (Mayer et al., 1985b; Lasic et al, 1992; Haran et al, 1993; Cheung et al, 1998; Fenske et al., 1998), and as indicated in the following section these active loading procedures played a fundamental role in the development of clinically viable anti-cancer drug formulations. 1.3 Liposomes as drug carriers Research on liposomes as model membrane systems and as drug carriers facilitated the design of pharmaceutically viable lipid-based drugs. In fact much of the research and technology required to prepare liposomal carriers for testing in clinical trials was well established by 1987 (Cullis et 15 Figure 1.4 Active drug entrapment in liposomes Redistribution of a lipophilic amine (weak base) in response to a pH gradient (ApH) across the liposome membrane. Only the neutral form of the molecule is capable of crossing the lipid bilayer. (Figure reproduced from Parr, 1995) Outside Inside pH 7.5 pH 4.0 K = [B].[H*]„ K = [B],[H+], [BH +] 0 [BH+1 At equilibrium, if: [B]0 = [B], Then: [BH+1 [H+], f 3 r? ] 0 Fio 16 al, 1987; Ostro and Cul l is , 1989; Perez-Soler, 1989). B y that time, four pivotal hurdles were overcome. First, the importance of carefully assessing structure activity relationships through analysis of physiochemical characteristics was proven to be essential in product development. This is exemplified by studies contributing to the characterization of the amphotericin-B lipid complex (Janoff et al, 1988; Grant et al, 1989). Second, biological barriers previously believed to limit the distribution properties of systemically administered macromolecular drug carriers, such as liposomes, proved to be penetrable. In 1983, John Baldeschwieler and co-workers recognized that liposomal drugs could effectively deliver contents to tumours (Proffitt et al, 1983), a phenomena that continues to be a fundamental rationale for development of systemically administered liposomal anti-cancer drugs (Gabizon and Martin, 1997). Third, manufacturing issues for preparing pharmaceutically acceptable formulations were resolved (Lichtenberg and Barenholz, 1988; Swenson et al, 1988; Vuil lemard, 1991). This included identification of sources for inexpensive raw materials, the elucidation of procedures for storing lipid-based carriers for extended time periods (Madden et al, 1985) and the development of methods for reproducibly preparing large batches of liposomes with attributes that could be characterized according to the rigorous guidelines of health boards such as the F D A . Fourth, procedures for loading liposomes with pharmaceutical ly active agents that relied on the chemical attributes o f the lipids prior to liposome formation (e.g. doxorubicin/cardiolipin complex) and/or involved loading of pre-formed liposomes were developed (Wizke and Bittman, 1984; Gootmaghtigh et al, 1987 Mayer et al, 1990; Schwendener et al, 1991; Haran et al, 1993). The latter involves the use of ion gradients to effect drug loading (see section 1.2.4.2), a procedure that has proven to be particularly useful and versatile. A t the end o f the 1980's investigators confidently suggested that liposomes could be designed to achieve specific therapeutic benefits for a broad range of disease targets. It is perhaps 17 disappointing, therefore, that improvements in the therapeutic properties of liposomal drugs have been relatively incremental since 1990. The most significant revisions of lipid-based carrier technology that have guided research efforts during the 1990's involved three breakthroughs made in the late 1980's: 1) the observation that surface associated polymers (i.e. polyethylene glycol or the ganglioside G M 1 ) cause changes in the liposome surface properties that contribute to increased circulation lifetimes (Allen and Chonn, 1987; Papahadjopoulos et al, 1991); 2) the discovery that positively charged liposomes can be used to transfer polynucleotides into cells (Brigham et al, 1989; Feigner and Ringold, 1989); and 3) the identification o f certain lipids that can act as therapeutic molecules (Berdel et al, 1986). 1.3.1 Liposomal anti-cancer drugs There are two general reasons for developing a liposomal anti-cancer drug. First, the drug may be hydrophobic and difficult to dissolve in aqueous solutions, and thus a hydrophobic environment is required in order for the drug to remain in solution/suspension. Second, the liposome can serve as a carrier that wi l l improve drug specificity by increasing delivery to the site of disease and/or decrease delivery to a site where toxicity is manifested. The former is an important, perhaps underdeveloped, role for lipid-based carriers. However, the methods and characterization studies required for development of lipid-based formulations optimal for drug solubilization are distinct from those used in the development of liposome drug carrier technology. Differences in the two approaches can be defined primarily through in vivo studies that determine plasma elimination behavior of both drug and liposomal l ipid. If the drug dissociates from the liposome immediately following administration then the lipid-based carrier is acting as an excipient for drug solubilization. When drug elimination parameters are dictated by the elimination behavior of the liposomes, then the systems are acting as true delivery vehicles. 18 This thesis focuses on use of liposomes developed as drug carriers. The primary consequence of anti-cancer drug encapsulation is liposome-mediated changes in drug elimination and biodistribution. It is important to recognize that therapeutic responses obtained following administration of anti-cancer drugs, in free form or associated with a drug carrier, are dependent on tumour physiology and tumour cell heterogeneity. Ideally, an effective drug must access the target cell populations at levels sufficient to cause cytotoxic effects and should be effective in all microenvironments present within tumours. In humans, strategies designed to maximize the anti-tumour activity of chemotherapeutic agents must, therefore, contend with a heterogeneous population of proliferating cells. Tumour cells are proliferating at different rates, are governed by differences in cell cycle control and are capable of adapting rapidly to the chemotherapeutic stresses exerted on them. In practical terms this means that chemotherapy typically involves the use of multiple drugs that exert anti-tumour activity via different mechanisms (De Vi ta , 1997). Vincristine is a cell cycle specific agent that acts by destabilizing microtubules and is almost always used in combination with two or three other anti-cancer drugs. The therapeutic action of vincristine is complemented by drugs such as doxorubicin (an anthracycline that acts as a topoisomerase II inhibitor) as well as cyclophosphamide (a nitrogen mustard pro-drug and strong alkylating agent). The mechanisms of therapeutic action of these drugs are quite different; this complementary nature and the side effects of each drug are sufficiently different such that they can be used in combination, thereby increasing the reduction in tumour burden and decreasing the risk of drug resistance. In addition to the necessity of using multiple agents to achieve optimal therapy, another general principle of cancer chemotherapy concerns maximizing dose intensity (Livingston, 1994). Tumour cells must be exposed to the highest levels of drug for the longest time periods i f maximum therapeutic effects are to be achieved (Mulder and de Wit , 1995). The advantage of anti-cancer drug carrier technology is based on carrier characteristics that give rise to increased 19 drug exposure in sites of tumour growth. A n example of how liposome drug carrier technology can improve the pharmacodynamic behavior of an anti-cancer agent is evident when evaluating studies with doxorubicin. Efforts to maximize the dose intensity of this chemotherapeutic agent (in free form) have been limited due to non-specific toxic side effects. Therapeutic doses must, therefore, be limited to schedules and amounts that do not compromise regeneration of blood cells or cells of the immune system. In addition, doxorubicin exhibits a dose limiting cardiotoxicity (Minow et al, 1975) restricting the total dose to approximately 450 mg/m 2 . Myelosuppression can be counteracted using the hemopoietic growth factor granulocyte-macrophage colony-stimulating factor ( G M - C S F ) (Elias et al, 1993). Administering the drug in a liposomally encapsulated form, on the other hand, can reduce cardiotoxicity (Gabizon et al, 1982; Herman et al, 1983; Balazsovits et al, 1989). It has also been shown that the therapeutic activity of the liposomal drug is greater than or equal to free doxorubicin in a variety of pre-clinical and clinical studies (Mayhew et al, 1987; Mayer et al, 1990; Elias et al, 1993; Northfel te /a / . , 1997; V a i l etal, 1997). Table 1.2 summarizes information on some of the major anti-cancer drugs that have been evaluated in a liposomal formulation. The formulations that have advanced the furthest along the clinical development pathway includes those used for doxorubicin (approved for clinical use in A I D S related Kaposi 's sarcoma), daunorubicin (approved for clinical use in A I D S related Kaposi ' s sarcoma); cisplatin (Phase I clinical trials; Perez-Soler et al, 1990), and mitoxantrone (Phase I/II clinical trials, Pestalozzi et al, 1995). The studies developed in this thesis have focused on the anti-cancer drug mitoxantrone. Some of the rationale for selecting this drug have been summarized below. In addition, data summarized in Chapters 3-5, add to these rationale and suggest that mitoxantrone is a excellent drug to consider for development as a liposomal formulation. 20 Table 1.2 M a j o r antineoplastic agents evaluated in a liposomal drug carrier system Class /Drug # of Different Liposomal Formulations Pre-clinical Evaluations Cl in ica l Testing Plant Alkal iods-Vincristine <10 extensive Phase II Vinblastine <5 very limited — Anthracyclines-Doxorubicin >10 extensive Approved Daunorubicin <5 extensive Approved Mitoxantrone <5 extensive Phase II Antimetabolites-Methotrexate <5 limited — 5-Fluorouracil <5 limited — Cytosine arabinoiside <5 limited — Other cis-diamminedichloroplatinum <5 limited — 1.3.2 Mitoxantrone Mitoxantrone is a dihydroxyanthracenedione (as shown in Figure 1.5) that was developed in an effort to produce new agents with similar modes of action to doxorubicin without the cardiotoxic side effects. It has demonstrated activity in a wide range of experimental tumours such as P388 and L1210 leukemias, A D J PC6 plasmacytoma, B16 melanoma, colon and mammary adenocarcinomas, transitional cell bladder carcinoma, and M5076 carcinoma (Johnson et al, 1979; Wallace etal, 1979; Corbett eta!., 1982; Fujimoto and Ogawa, 1982; Schabel et al., 1983 a,b; Bal lou and Tseng, 1986). It has been investigated in the treatment of advanced breast cancer (Brambilla el al, 1989; Harris et al, 1990), non-Hodgkin's lymphoma (Bajett et al, 1988, Ho et al., 1990), acute leukemia (Bezwoda et al., 1990; Amadori et al, 1991; Archimbaud et al, 1991; Hiddemann et al, 1991; Wahlin et al, 1991), and hepatocellular carcinoma (Dunk et al, 1985; 21 Yoshida et al, 1988; Lai et al, 1989; Colleoni et al, 1992). Mitoxantrone has proved useful as palliative therapy in patients with hepatocellular carcinoma (Civalleri et al, 1996), advanced breast cancer (Roberston et al, 1989), or prostate cancer (Tannock et al, 1996). Figure 1.5 Structure of mitoxantrone hydrochloride There are several mechanisms of action that have been identified: 1) D N A intercalation, 2) stabilization of the topoisomerase-DNA complex, 3) D N A condensation via electrostatic cross-linking, and 4) non-protein-associated D N A strand breaks induced by free radical generation via oxidative activation. Structurally mitoxantrone is similar to doxorubicin in that it contains a planar polycyclic aromatic ring structure which allows it to intercalate within the D N A and inhibit D N A and R N A synthesis (Safa et al, 1983; Durr, 1984). In addition, mitoxantrone induces protein associated strand breaks via stabilization of the topoisomerase II complex. Topoisomerase enzymes are responsible for the catalysis of the breaking and rejoining of D N A via an enzyme-DNA intermediate, topoisomerase I for single strands and topoisomerase II for 22 double strand breaks. Mitoxantrone appears to inhibit topoisomerase II by binding to the enzyme-DNA complex, thereby preventing rejoining of the D N A . Unlike doxorubicin, which is reduced to a semiquinone free radical via N A D P H cytochrome P-450 reductase, mitoxantrone does not produce any free radicals by this pathway and acts as a potent antioxidant (Fisher et al, 1989; V i l e and Winterbourn, 1989). However, mitoxantrone can undergo peroxidative conversion to an unstable diimino compound which then generates a radical cation. This oxidative activation results in D N A damage as demonstrated by Fisher and Patterson (1989). Several liposomal formulations of mitoxantrone have been developed for the treatment of cancer. Two groups have focused on passive entrapment of mitoxantrone in liposomes (Schwendener et al, 1991; L a w et al, 1996). Phase II clinical trials have been conducted but demonstrated disappointing activity in the treatment of breast cancer. The formulation tested, however, exhibits rapid release characteristics and /or liposome elimination with the majority of the liposomes eliminated from the plasma compartment within the first 10 minutes (Schwendener et al, 1994). Thus, the benefits of using a liposomal carrier were not observed. In addition, the use of charged liposomes increased the accumulation of the liposomes to the liver and spleen. Other formulations developed have utilized the pH gradient encapsulation of mitoxantrone (Madden et al, 1990; Schwendener et al, 1994; Chang et al, 1997). These formulations demonstrate significant increases in drug circulation lifetime and levels of drug within the plasma compartment when compared to the free drug (Schwendener et al, 1994; Chang et al, 1997;), leading to improvements in efficacy over the free drug (Chang el al, 1997; L i m et al, 1997). 23 1.4 Biological fate of liposomes following intravenous administration In vivo studies are usually initiated only after the development of a liposomal formulation that exhibits the necessary chemical and physical stability properties to be considered pharmaceutically viable. A s suggested in section 1.3, technological advances as well as an increased understanding of lipid chemistry have, to large extent, overcome many pharmaceutical hurdles. This section w i l l focus on systemic administration and, in particular, on the fate of lipid-based delivery systems injected intravenously (i.v.). 1.4.1 Barriers and compartments In vivo analysis must consider the fact that a liposomal drug wi l l interact with a number of distinct physiological "compartments" and associated barriers between compartments. After injection, liposomes are exposed to a variety of circulating protein and cellular components that reside within the central blood compartment, many of which can destabilize the liposomes through interactions with the lipid bilayer or initiate biological processes that lead to increased liposome leakage and/or clearance via the mononuclear phagocyte systems (Allen and Cleland, 1980; Bronte et al, 1986; L iu et al, 1997). To gain access to a disease site in an extravascular compartment, liposomes must cross the vascular endothelium, the blood vessel lining which is composed primarily of endothelial cells and, in most cases, an underlying basement membrane and associated smooth muscle cells (See Figure 1.6). This vascular barrier represents the greatest obstacle for liposomal drug delivery to extravascular disease sites, however, at the same time it offers properties that can be utilized to differentiate between normal and diseased tissue. Should liposomes traverse this barrier, a second compartment is encountered consisting of the interstitial space and associated fluids and cells. This compartment can vary significantly not only between normal and disease tissues but also among normal tissues in different organs o f the 24 Figure 1.6 Structure of a capillary: continuous endothelium type Electron micrograph of a capillary composed of continuous endothelium. The capillary is supported by a basement membrane ( B M ) and collagen fibrils (C). A pericyte (P) embraces the capillary and is supported by its own basement membrane (BMp) . The endothelial cells are seen encircling the capillary lumen with cytoplasmic flaps called marginal folds (M) extending across the intercellular junctions. (Reproduced from Burkitt et al., 1993) 25 body. Within this compartment, the barriers to liposome movement and distribution are varied and include factors such as interstitial volume, interstitial pressure, and the presence (or absence) of a lymphatic system. The final physiological compartment(s) is the cells into which liposomes and/or their associated agents are taken up. This includes intracellular organelles that may be involved in processing o f the administered agent or that contain the molecular target through which the drug exerts its therapeutic activity. The critical barrier that must be crossed in order to access this final compartment is the cell membrane. Similar to the vascular endothelium, crossing this barrier is a significant obstacle to the development of therapeutically optimized liposomal anti-cancer drugs. In the following sections, the fate of liposomes w i l l be discussed as they enter these physiological compartments and pass through the various barriers. The focus w i l l be on specific interactions between liposomes and the biological milieu in the various compartments that directly impact on the delivery of encapsulated agents to their therapeutic target. Further, sections w i l l highlight strategies that have been employed to augment conventional liposomes (defined as un-derivatized membrane bilayers composed of naturally occurring lipids) with components that alter these interactions. 1.4.2 Liposome serum protein interactions Within the blood vessels, liposomes are exposed to circulating cells, lipoproteins, other serum proteins as well as other small molecules such as carbohydrates and divalent cations. A s indicated in section 1.2.2, liposomes designed for intravenous application typically contain 30 to 50 mol % cholesterol, a required component to minimize the protein-liposome interactions (Patel et al, 1983; Semple et al., 1996). It is important to recognize that cholesterol-containing liposomes bind other serum proteins (Bonte and Juliano, 1986; Chonn et al, 1992) and the 26 biological fate o f the liposome is determined, in part, by these associated proteins. Serum protein binding can increase membrane permeability as well as play a role in defining the liposome elimination rate and biodistribution characteristics. These two effects are discussed in the following sections. 1.4.2.1 Serum-induced increases in membrane permeability Serum protein binding can increase liposome permeability non-specifically and specifically. The latter is best illustrated by studies assessing liposome-complement protein interactions (Silversmith and Nelsestuen, 1986a, b, Mal inski and Nelsestuen, 1989, Shiver et al, 1991). It is known, for example, that anionic liposomes (those containing PS or cardiolipin) can activate the alternative complement pathway that is associated with C3b binding and formation of the membrane attack complex ( M A C ) . The M A C is a complex of the complement proteins C5b-6, C7 , C8 and C9 and its formation has been associated with increased membrane permeability (Mal inski and Nelsestuen, 1989) attributed to ion channel formation and/or pore formation as wel l as transbilayer flip-flop of lipids (Van der Meer et al. 1989). In addition, complement binding has been shown to influence the binding of C-reactive protein (CRP) ( L i el al, 1994) and vise-versa (Richards et al, 1977, Richards et al. 1979). C R P is known to localize in sites of inflammation or other regions displaying membrane damage and it plays a role in recruitment of inflammatory cells including those phagocytic cells involved in removal of damaged cells. Binding of C R P w i l l , therefore, play a role in immune recognition of certain liposome formulations. This is typically not an important factor when the liposomal formulations being developed are composed of neutral lipids, such as PC and cholesterol. In the context o f this thesis, non-specific effects of serum protein binding on liposome permeability are those that can not be attributed to a defined protein or complex o f proteins. 27 Serum-induced increased release of encapsulated drugs (e.g. vincristine) and markers (e.g. calcein) (Al len and Cleland, 1980; Boman et al. 1993) can be determined in vitro and in vivo. Interestingly, release rates measured in vitro in the presence of serum are often much slower than those measured following i.v. administration. The in vivo data are determined by monitoring changes in the drug to lipid ratio o f liposomes within the plasma compartment (Figure 1.7). A s shown in Chapter 3, drug leakage rates can be significantly greater in vivo than in vitro. This is consistent with other reports that stress that in vivo drug retention properties and drug release kinetics for different liposomal formulations can not be predicted on the basis of in vitro data (Bally et al, 1993). Another serum protein mediated effect on membrane permeability is particularly unique to the active loading procedure, such as that used in this thesis. The high concentrations o f buffer components and/or entrapped drug in liposomes can result in a significant osmotic gradient across the liposome membrane when exposed to physiological fluids. It has been shown that liposomes can withstand a transmembrane osmotic gradient of greater than 100 mOsm/kg in the absence of serum proteins; however, these liposomes release a portion of their contents when diluted into serum containing buffers ( M u i et al, 1994). This is typically seen as a burst o f entrapped-content release that occurs while an osmotic balance across the membrane is re-established. 1.4.2.2 Serum protein binding and liposome elimination In a general context, there appears to be a direct correlation of increased protein binding to liposomes and increased liposome elimination rates (Chonn et al, 1992). Increased protein binding and clearance are, in particular, identified with liposomes composed of anionic lipids (e.g. phosphatidylserine, cardiolipin and P A ) (Spanjer et al, 1986; Chonn et al, 1992) and cationic lipids (e.g. stearylamine) (Mold et al, 1981; Oku et al, 1996). Certain proteins such as complement proteins, serum albumin and beta 2 glycoprotein 1 have been associated with 28 Figure 1.7 Liposomal drug release Drug release from liposomes in vivo can be estimated by measuring drug-to-lipid ratio of liposomes in the blood compartment. It is important to recognize that two events are being monitored as a function of time after i.v. administration. Liposomes are being eliminated from the plasma compartment and, in addition, drug is being released from the liposomes. 29 increased elimination rates and these have, in turn, been attributed to their role as opsonizing proteins that are instrumental in "labeling" foreign macromolecules in the blood compartment (Chonn et al, 1995). This is an essential component of the immune system that facilitates recognition o f bacteria and damaged/dead cells by cells of the mononuclear phagocytic system ( M P S ) (see section 1.4.4). It is notable that not all anionic lipids cause an increase in liposome clearance. Phosphatidylglycerol and phosphatidylinositol containing liposomes exhibit plasma elimination rates that are comparable to or slower than neutral liposomes and this is in spite of having increased levels of absorbed serum proteins (Chonn et al, 1992). It is also worth noting that protein binding can have a direct effect or an indirect effect on liposome elimination from the plasma compartment. The indirect effect is one related to "opsonization" of liposomes and subsequent recognition by the M P S . PS containing liposomes, for example, are eliminated rapidly following i.v. administration due to this opsonization effect (Spanjer et al, 1986). In contrast, P G containing liposomes (when administered to rats) bind the complement protein C3b which is subsequently converted to C3bi . The presence of bound C3bi facilitates liposome binding to platelets that express the C3bi receptor, and an associated platelet aggregation reaction occurs (Reinish et al, 1988; Doerschuk et al, 1989). This aggregation reaction leads to removal of the aggregates in certain vascular beds such as those in the lung and spleen. This elimination mechanism is not affected by the M P S . It has also been postulated that saturation of serum protein binding can occur, i.e. that there is a limited amount of blood protein that is available to bind to liposomes (Oja et al., 1996). It has been demonstrated that at liposome doses ranging from 10 to 100 mg lipid/kg animal weight, circulation lifetime increases. In addition, the amount of protein bound to the liposome decreases as the dose increases. However, it is believed that at higher doses saturation o f the M P S system occurs, resulting in the increased circulation lifetime (Abra and Hunt, 1981; Bosworth and Hunt, 1982). To date the exact mechanism of liposome elimination has yet to be 30 elucidated but two factors affecting liposome elimination kinetics are: 1) the role of the M P S system and 2) the degree of protein binding. 1.4.3 Extravasation through blood vessels A microvascular structure is a capillary network composed of endothelial cells, a basement membrane, connective tissue elements, associated marginated leukocytes and the presence of certain serum proteins which function collectively as a selective barrier to circulating cells and macromolecules. In addition, the microvascular structure serves to selectively determine what size macromolecules can penetrate the blood vessel and this in turn is dependent on the tissue type and/or the presence of disease (Dvorak et al, 1988). With regards to liposomal drug carriers systems there is compelling evidence (both theoretical and experimental) that these circulating macromolecules wi l l have limited access to extravascular sites (Jain and Baxter, 1988; Yuan et al, 1995). Liposomes with a mean diameter in excess o f 50 nm w i l l only leave the blood compartment in tissues where large pores or fenestrations exist in the associated blood vessels. Blood vessels of the liver and spleen provide examples o f such tissues. However, there is also substantial evidence, albeit phenomenological, that liposomes can access extravascular sites within tumours following intravenous administration (Gabizon A . A . , 1988; 1992; Yuan et al, 1994). 1.4.3.1 Tumour vasculature It is established that tumours can exhibit unique microvascular structures that are often incapable of maintaining a complete permeability barrier between the vascular compartment and the growing tumour mass (Heuser and Mi l le r , 1986; Dvorak et al, 1988). Thus there are potential 31 sites where large drug carriers can escape from the circulation. Blood vessels of particular interest include; 1) sinusoidal vessels which are extremely porous, exhibiting large gaps between endothelial cells that are not closed by any membrane structure providing a discontinuous endothelium; 2) capillaries which exhibit fenestrated endothelium characterized by pores between endothelial cells, which allow macromolecules in the range of 20 to 100 nm to pass; 3) blood channels which lack an endothelial cell lining, allowing blood to percolate around and between tumour cells; and 4) postcapillary venules in tumours which are characterized by vessel walls composed of endothelial cells, devoid of basement membrane, supported by some fibrous tissue. The presence of these blood vessels in tumours w i l l promote leakage of circulating liposomes. A s indicated above, the organization endothelial cells adopt in different blood vessels influences the permeability characteristics of blood vessels. Endothelial cells also participate more directly in normal physiological processes regulating microvascular permeability (Simionescu, 1983; Simionescu et al, 1987; Pearson, 1991; Crone, 1986). These cells, for example, are known to have a direct role in the transport of serum components to extravascular compartments. Proteins and other circulating macromolecules can be taken up by endothelial cells via receptor mediated and fluid phase endocytosis (Simionescu, 1983; Simionescu et al, 1987). Subsequently the internalized material can be either degraded by transfer to lysosomal compartments (Ryan, 1988) or alternatively the endosome contents can be moved through the cell and released into the interstitial space on the opposite side of the cell (Kohn et al, 1992; Dvorak et al, 1996). This latter process is referred to as transcytosis. Further, it is known that endothelial cells are capable of phagocytosis and can actively accumulate particles in excess of 5 um in diameter (Ryan, 1988). Given these characteristics it is reasonable to postulate that endothelial cells play an important role in governing the fate of liposomal drug carriers. 32 In addition, factors secreted by the tumour associated cells also affect the permeability of the vasculature. The most dominant factor is vascular endothelial growth factor ( V E G F ) . This protein has been associated with several characteristics of tumour blood vessels such as increased vascular permeability (Dvorak et al, 1991), increased transcytotic activity and angiogenesis (Folkman and Shing, 1992). The endothelial cells express a high-affinity receptor for V E G F (De Vries et al, 1992; Takagi et al, 1996) which is a member of the platelet-derived growth factor receptor family, fit. Expression of these high-affinity receptors can be induced in tumour vascular endothelial cells (Senger et al, 1993) and in endothelial cells maintained under hypoxic conditions (Stein et al, 1995). 1.4.4 Role o f the mononuclear phagocytic system (MPS) The M P S [previously referred to as the reticuloendothelial system (RES)] has long been recognized as the major site of liposome accumulation after systemic administration. The primary organs associated with the M P S are the liver, spleen and lung. The liver exhibits the largest capacity for liposome uptake while the spleen can accumulate liposomes such that the tissue concentration (liposomal lipid/g tissue) is as much as 10-fold higher than that which can be achieved in other tissues. Assuming that liposomes are designed to minimize protein binding (see section 1.4.2.2) and cell interactions, the extent of liposome accumulation in the lung is typically below 1% of the injected dose. Early studies demonstrated that large, as well as charged liposomes (particularly those containing negatively charged lipids like PS, P A or cardiolipin), were removed very rapidly by the liver and spleen with more then 50% of the injected liposomes being eliminated from the plasma compartment in less than 1 hour (Chonn et al, 1992). However, when small (approx. 100 nm), neutral liposomes containing > 30% cholesterol are injected at doses of at least 10 mg/kg or more the plasma elimination rate is 33 substantially reduced (Patel et al, 1983; Semple et al, 1996;). The removal of liposomes from the blood is attributed to phagocytic cells that comprise the M P S and uptake of liposomes by cells of the M P S is mediated through direct interactions between the phagocytic cell and the liposomes and is stimulated by the binding of certain serum proteins (Chonn et al, 1992). When the dose of the liposomes is increased to levels of 100 mg/kg, there is a further increase in the circulation longevity of the liposome carrier. This is due to two effects: saturation of the M P S (Abra et al., 1981; Bosworth et al., 1982) and depletion of circulating opsonins which mark the liposomes for elimination (Oja et al., 1996). If liposomes are designed in an appropriate manner, whether with respect to size or lipid composition, liposomes can remain in the blood compartment for a period of several days (Parr et al, 1997). The fact that under such circumstances the vast majority of liposomes administered can be accounted for in the blood, liver and spleen demonstrates that liposomes are relatively inefficient at crossing the endothelial cell barrier present in most other normal tissues. 1.4.4.1 Liposome accumulation in the liver The liver represents a major obstacle for liposomal formulations that are being designed for extravascular sites such as tumours residing in sites located away from the liver. Liposomes rapidly accumulate in this organ due to 1) the blood supply and vessel structure and 2) the presence of Kupffer cells. The liver is unique in that it has a dual blood supply from the hepatic artery and the portal vein. Therefore, any intravenous injection w i l l pass through the liver. Further, the liver functions as a filter, removing unwanted debris (e.g. senescent erythrocytes, bacteria, and toxins), as well as a detoxification organ. The cells responsible for this filtration process are the Kupffer cells which phagocytose and remove any foreign elements. Thus, following /.v. injection of a liposomal formulation, the liposomes wi l l naturally accumulate in the liver due to the blood supply and be processed by the Kupffer cells. 34 There have been many studies investigating liposomal interactions with the cells of the liver including hepatocytes, endothelial cells and Kupffer cells (Hu and L i u , 1996; Spanjer et al, 1986; Kamps et al, 1997;). The architecture of the liver is such that the blood percolates through sinusoids lined by endothelial cells and Kupffer cells. The endothelial layer which lines the sinusoids is discontinuous allowing macromolecules, ranging in size from 70 nm to 120 nm, to access to the hepatocytes. These molecules then enter the space of Disse between the endothelium and hepatocytes allowing for interaction with liver cells outside the blood compartment (see Figure 1.8). Hepatocytes are organized into one or two-cell-thick plates that are separated by sinusoids. They are responsible for the major functions of the liver such as detoxification, protein synthesis, metabolism, and storage. Numerous studies have been performed to understand the role of liver in liposome clearance and many pathways have been postulated for liposome uptake in the cells of the liver (Hu and L i u , 1996; Scherphof and Kamps, 1998). Many of these pathways involve receptor mediated endocytosis (Scherphof and Kamps, 1998), or serum protein binding (Hu and L i u , 1996). A s indicated in section 1.4.2.2, complement proteins and A p o E have been implicated in the removal of liposomes from the circulation (Chonn et al, 1995; Devine and Bradley, 1998; Scherphof and Kamps, 1998) and liposomes that exhibit a negative charge, such as phosphatidylserine containing liposomes, are rapidly taken up by the Kupffer cells due to the increased amount of protein bound to these liposomes. In general, neutral liposomes composed of P C with a mean diameter of 100 to 200 nm w i l l also localize in the Kupffer cells; however, the rate at which these liposomes accumulate in this cell population is much slower. A n y population of liposomes that exhibit a diameter of less than 50 nm do, however, have the potential to interact with the hepatocyte population in the liver (Scherphof et al, 1987). The interaction of liposomes with 35 Figure 1.8 Diagram of a classic liver lobule Branches of the hepatic artery ( H A ) and hepatic portal vein (PV) empty blood into hepatic sinusoids (S), through which it flows toward the central vein. The endothelial lining of the sinusoids is discontinuous and is separated from the radial plates of hepatocytes by the space of Disse. B i l e canaliculi receive bile from the hepatocytes that border them and convey it toward the bile ducts in the portal triads. The arrows show that blood (dark arrows) and bile (open arrows) flow in opposite directions. (Figure reproduced from Paulsen, 1996) 36 Kupffer cells and hepatocytes and the relationship with therapeutic activity is still not well understood as demonstrated by the data presented in Chapter 5. 1.4.4.2 Decreasing liposome interactions with the M P S The identification o f certain naturally occurring lipids (e.g. ganglioside G M i and PI) (Al len and Chonn, 1987) and synthetic lipids with selected polymers linked to the head group (Al len et al, 1991; Yuda et al, 1996;) that decrease the plasma elimination rate of liposomes has provided a fundamental advance in liposome technology. It is believed that these lipids act by limiting the interaction of liposome surfaces with proteins and this, in turn, inhibited the rate of uptake by phagocytic cells (Chonn et al, 1991, 1992). The best characterized example of these lipids is based on the hydrophilic polymers P E G which can be chemically linked to the reactive amine function of the P E . The steric stabilizing lipid that is used most frequently is composed o f 2,000 mean molecular weight linear P E G moiety attached to D S P E . This lipid is incorporated into the liposomes while being prepared, typically at levels less than 10 mol %. Inclusion of P E G - P E into neutral (PC/cholesterol) liposomes can result in 3 to 20-fold increases in plasma liposome content 24 hour after i.v. injection (Allen et al, 1991; Parr et al, 1997). This is accompanied by significant decreases in liposome uptake by the liver and spleen at early times post-injection. It is important to note that the difference in cumulative uptake by the liver and spleen of liposomes with and without P E G - P E are reduced as a function of time, indicating that the effect of P E G - P E is to reduce the rate of liposome removal by cells of the M P S . It has been demonstrated recently that PEG-modified lipids can be lost from the outer monolayer of the liposomal membrane due to lipid transfer or cleavage of the PEG-linker and it is not clear whether eventual removal o f P E G liposomes by the M P S is due, in part, to the loss of P E G moiety (Parr et al, 1994). 37 Significant increases in circulating levels of liposomes can also be achieved by strategies that eliminate phagocytic cells of the M P S . This effect, referred to as M P S (RES) "blockade", can be achieved by pre-dosing animals with a low dose (10 mg lipid/kg) of liposomal doxorubicin (Bally et al, 1990; Daemen et al, 1995,) or alternatively through use of the encapsulated bisphosphonate clodronate (Van Rooijen and Claassen, 1989; Van Rooijen et al, 1990) (see Chapter 5). Investigators have been able to demonstrate macrophage and Kupffer cell depletion following administration of high doses of large and/or negatively charged liposomes containing doxorubicin or other agents such as clodronate (Van Rooijen et al, 1990; Daemen et al, 1995). M P S blockade induced by low doses (<10 mg/kg lipid and 2 mg/kg drug) of small, uncharged liposomal doxorubicin formulations, however, does not result in complete elimination o f Kupffer cells (see Chapter 5). The M P S blockade effect observed for liposomal anti-cancer drugs has raised concerns over potential harmful side effects resulting from altered phagocytic cell activity. Although a substantial amount of doxorubicin can accumulate in liver tissue, indications of significant liver toxicity arising from this uptake have only been observed pre-clinically with high drug doses (80 mg doxorubicin/kg) and in clinical situations where pre-existing liver impairment was a factor. It should also be stressed that the theoretical "benefits" arising from decreased liposome elimination by the M P S is typically assumed to be related to the increased circulating concentrations of liposomes obtained. For example it has been suggested that maintenance of the plasma concentration of liposomes for extended time periods is essential to maximize the amount of liposomal drug that penetrates the vascular barrier and gains access to diseased tissue. In this thesis M P S blockade is used to address the importance o f liver phagocytic cells in mediating the therapeutic activity of a liposomal formulation of mitoxantrone (see Chapter 5). 38 1.4.5 Liposome extravasation A s discussed in section 1.4.3, diseases such as bacterial infection, inflammation and cancer share a common feature in that the diseases induce regional increases in vasculature permeability. The mediators that lead to increased permeability of the vascular barrier are quite distinct and can be attributed to transendothelium migration of inflammatory cells (Thureson-Klein et al, 1986; K l i n g et al, 1987) or to the release of vascular endothelial growth factor ( V E G F ) (Senger et al, 1993). Regardless of the mediator, the end result for all of these conditions is the presence of blood vessels that are permeable to large molecules. This may be a consequence o f fenestrations or "gaps" occurring between adjacent endothelial cells through which macromolecules can pass (Jain, 1987). Alternatively, liposome extravasation may involve increases in endothelial cell mediated transcytosis (Kohn et al, 1992; Dvorak et al, 1996). Increases in vascular permeability give rise to the selective accumulation of small liposomes at sites of infection, inflammation and tumour growth. However, this is not a selective process. There is also a general increase in extravascular fluids in these regions. The hydrostatic pressure within these sites is elevated relative to the vascular pressure, resulting in a pressure gradient that impedes movement of molecules from the blood into the tissue interstitium (Baxter and Jain, 1989). It must therefore be assumed that additional features lead to selective accumulation of macromolecules in the diseased extravascular space. Studies, for example, have demonstrated that the lack of a developed lymphatic system in conjunction with the large openings in the vascular endothelial cell lining may lead to an extravascular "trapping" phenomenon (Baxter and Jain, 1990). In the absence of lymphatic drainage, interstitial diffusion of molecules leads to egress from the disease site and this diffusion rate is dependent on molecule size, small molecules exiting more rapidly than large molecules. 39 Designing liposomes that wi l l exhibit maximal extravasation in disease sites associated with leaky vasculature is of considerable interest and is an area of some controversy. The inclusion of PEG-modif ied lipids in conventional liposomes can significantly increase the circulating liposome levels over extended times by decreasing the rate of clearance by the M P S . It has generally been assumed that increases in the concentration of liposomes in plasma over time w i l l lead to increased accumulation of liposomes in the extravascular disease sites, and experimental evidence supporting this has been reported (Gabizon, 1992). However, there are studies contrasting these reports. It has been demonstrated that although plasma levels o f P E G containing liposomes are several fold higher than for comparable conventional liposomes, this often does not result in increased extravasation and accumulation in solid tumour tissue (Parr et al, 1997). It should not be unexpected that conventional and sterically stabilized liposomes exhibit different efficiencies in extravasation. Endothelial cell interactions may contribute to the extravasation process either directly via transcytosis or indirectly by facilitating an increase in the local liposome concentration at the endothelial cell surface. Given the effects of P E G on inhibiting liposome-cell interactions (Du et al, 1997), this polymer may reduce endothelial cell interactions and this, in turn, would reduce the rate of extravasation. 1.5 Dissociation of the active agent from the carrier: the critical parameter The distribution of liposomes that have extravasated into the tumour interstitium is heterogeneous and these large carriers diffuse slowly within the perivascular spaces (Yuan et al, 1994). Slow diffusion within the site of extravasation has also been associated with very slow loss of the liposomes from the site. Data from several tumour models, including results shown in 40 Chapter 4, demonstrate that the level of liposomes achieved following extravastion can be maintained for extended time periods (Parr et al, 1997). Importantly, drug accumulation properties in solid tumours or within other disease sites can exhibit remarkably different behavior in comparison to the liposomal carrier. Drug release from the liposomes in the extravascular site can result in greater drug penetration into the tissue and more rapid loss of the drug from the site when compared with the loss of liposomal lipid (see Chapter 4 for example). It is not clear from studies correlating anti-cancer activity and increased liposome mediated drug delivery, what is the critical parameter to consider when optimizing a liposomal anti-cancer drug. This thesis has the primary goal of addressing this problem. Studies demonstrating improvement in liposomal anti-cancer drug activity in comparison to free drug have typically compared the efficacy and drug accumulation following administration of drug doses that are equivalent on a weight basis (equal mg/kg dose) or toxicity basis (at the maximum tolerated dose). Under these conditions, there can be 3- to 100-fold increases in drug exposure achieved for the liposomal formulations. It is anticipated, however, that efficacy measured under conditions where tumour drug accumulation levels are comparable for free and liposomal drug that the liposomal drug would be less active. This assumption is made on the basis of studies that demonstrate significant (100-fold) increases in drug exposure, but only marginal (20%) increases in therapeutic activity (Parr et al, 1997). Such observations have raised obvious questions about the biological availability of anti-cancer drugs carried inside liposomes that have extravasated into solid tumours as well as the mechanisms that lead to drug release in the interstitial compartment. It can be suggested that liposomes exert their effect on the therapeutic activity of an associated anti-cancer drug by providing a drug infusion reservoir within the tumour. Once released, the anti-cancer drug can diffuse through the tumour, directly accessing tumour cells in a manner that 41 is comparable to drug in the absence of a liposomal carrier. There are questions regarding how and where drug release occurs and, as suggested in Chapters 4 and 5, a model consisting of drug release from liposomes in the plasma compartment or from a site distant from the disease can not be discounted. In vitro studies, for example, have demonstrated that macrophages can engulf doxorubicin-loaded liposomes, process them and re-release doxorubicin extracellularly in free form (Storm et al, 1988). Since the macrophage content within tumours can be significant, it can be suggested that liposomal anti-cancer drug release may involve macrophage processing after extravasation. Interestingly, however, recent studies have shown that interactions between tumour-associated macrophages and extravasated liposomes are minimal (Mayer et al, 1997). 1.5.1 Drug release - importance of drug type As indicated in section 1.4.2.1, it is not possible to predict drug release rates in vivo on the basis of in vitro studies even when the in vitro release studies are completed in the presence of serum. It is also not suitable to determine release rates using a trapped "marker" (e.g. radiolabeled inulin) to predict the release characteristics for an encapsulated therapeutic agent (Bally et al, 1993). Drug release rates are dependent on the chemical properties of the entrapped drug. This is perhaps best illustrated using liposomal formulations of vincristine and doxorubicin as described below. Reducing the drug release rate is advantageous for encapsulated formulations of vincristine but is of questionable benefit for doxorubicin. Liposome encapsulation can significantly reduce the toxicity of doxorubicin by decreasing drug accumulation in drug sensitive normal tissue, presumably by decreasing peak levels of free doxorubicin that are experienced after administration in the conventional (unencapsulated) form (Mayer et al, 1994). The degree of toxicity buffering is directly related to the ability of the liposomes to retain their entrapped 42 doxorubicin where increased phospholipid acyl chain saturation results in decreased toxicity (Mayer et al, 1994). The anti-tumour activity of liposomal doxorubicin, however, is much less sensitive to drug leakage or circulation longevity. Liposomal formulations with widely varying doxorubicin retention properties have been shown in some preclinical models to exhibit comparable anti-tumour activities when compared on an equal dose basis (Mayer et al, 1994). In this case, increased efficacy for the less permeable liposomes is achieved by administering elevated drug doses due to their reduced toxicity. Further, while the inclusion of P E G - P E increases the circulation longevity of liposomal doxorubicin, the magnitude of increased liposome levels in the blood (compared to conventional liposomes) is far less than that observed for empty (drug-free) liposomes (Parr et al, 1997). This is related to the M P S blockade effect described in section 1.4.4.2. In contrast to the observations made with doxorubicin, altering the physical properties of liposomal vincristine formulations results in dramatic changes in anti-tumour activity while only minimally affecting drug toxicity characteristics. Increasing the retention of vincristine inside 100 nm liposomes by changing the phosphatidylcholine-containing lipid component from E P C to D S P C to sphingomyelin (while maintaining cholesterol content at 45 mol%) leads to dramatic increases in anti-tumour activity, particularly when compared to the efficacy obtained with free vincristine (Webb et al, 1995). This is consistent with the steep dependence of vincristine anti-tumour potency on the duration of drug exposure as well as the fact that retention of vincristine in most tissues, including tumours, is rather poor. It appears that the ability to prolong the exposure of vincristine in vivo is more important than peak drug concentrations. Furthermore, although inclusion of P E G - P E in the liposomes increases the circulating liposomal lipid levels at extended time periods, this steric stabilizing lipid does not improve the vincristine pharmacokinetic or therapeutic properties over conventional D S P C / C h o l or sphingomyelin/Chol systems (Webb et al, 1998). This is due to the fact that P E G - P E increases the permeability of the lipid bilayer to vincristine, thus offsetting the potential benefits provided by increased 43 longevity of the liposomal carrier. It should be noted that perhaps the best example of how a balance between efficient liposome delivery to the disease site and controlled drug release can work synergistically to achieve optimum therapeutic results is provided by.the liposomal mitoxantrone data presented in this thesis. 1.5.2 Future considerations for the next generation of liposomes It can be suggested that-the drug retention properties required to minimize systemic exposure o f drugs encapsulated inside long circulating liposomes significantly limits biological availability of the agent once it has reached the disease site. This conclusion arises from results in several model systems that show that significant increases in disease site drug delivery often translate into only incremental increases in drug potency. It has been demonstrated in pharmacodynamic studies with liposomal anti-cancer agents that the circulating drug pool itself has little direct impact on therapeutic activity (Mayer et al., 1994). Instead, it appears that once extravasated, the lipid carrier provides a localized source of drug infusion within the disease site. While the liposomal drug formulations used to date have given rise to significant improvements in therapeutic activity, many results suggest that drug within the tumour is not freely biologically available. In vitro studies measuring the doxorubicin concentrations necessary for 50% inhibition of growth ( IC 5 0 ) of tumour cells in culture indicate a range in doxorubicin I C 5 0 ' s of 100 n M in M C F - 7 breast tumour cell line (Formari et al., 1994) to 190 n M and 24 u M in parental and D O X - resistant P388 cells, respectively (De Jong et al., 1995). Parr et al. (1997) has demonstrated that drug concentrations of 250 nmoles per gram tumour can be achieved using doxorubicin loaded drug liposomes and it can be suggested that drug concentrations within the tumour are in excess of that required to achieve maximum cytotoxic effects, even for drug resistant tumours. However, calculated rates of drug release from liposomes in tumour (0.60 to 44 0.65 nmol drug/umol lipid/h for doxorubicin encapsulated in D S P C / C h o l liposomes) may not be sufficient for inhibition or elimination of the tumour cells (Parr et al, 1997). The inability to differentially control drug release rates in the plasma compartment and disease site is perhaps the most significant limitation of presently available liposomes. A s suggested in section 1.1, it would be ideal i f one could design liposomes that have little or no drug leakage in the circulation and increased release rate at the disease site. Early attempts to selectively increase drug leakage at tumour sites centered on the fact that liposomes can be constructed to become leaky in the acidic interstitial p H of some solid tumours (Connor et al, 1984; Aicher et al, 1994), which can drop to values of 6.5. More direct evidence of the importance o f site-specific drug release has been obtained using localized hyperthermia (Chelvi et al, 1995; Gaber et al, 1996; Kakinuma et al, 1996). Liposomal doxorubicin preparations, for example, can be prepared such that there is an increase in drug release at 42°C, compared to 37°C. These liposomes are administered i.v. to tumour bearing mice and the tumour site is then heated using a topical microwave heating device placed on the subcutaneous tumour. Application of a transient heating pulse after the liposomal doxorubicin had accumulated into the solid tumour resulted in a significant increase of therapeutic activity compared to free drug with hyperthermia and liposomal doxorubicin in the absence of heating. Although hyperthermia may not be applicable to many multifocal or deep-seated tumours, this technique provides encouraging indications that liposomes exhibiting controlled or triggered release of their contents w i l l significantly augment the pharmacological improvements provided by liposomes. 1.5.3 Other methodology considerations For many applications, liposomal delivery systems are employed to improve the therapeutic index o f encapsulated agents by selectively accumulating in extravascular disease sites. A s 45 suggested above, there is also evidence indicating that drug released from liposomes in the circulation does not contribute significantly to therapeutic activity of liposomal anti-cancer agents. There is no question that liposomes can provide sustained exposure of therapeutic agents in the blood compartment through controlled release kinetics of encapsulated drugs; however, it is difficult to justify development of liposomal drugs using a rationale that involves sustained systemic exposure. This is largely due to significant advances made in the area o f drug infusion technology. Compact and cost effective infusion pumps are now widely used and these can provide well-controlled systemic drug exposure over several days. It is argued that the most significant advantage for the use of liposome drug carriers arises as a consequence o f disease specific changes in vascular permeability that favor accumulation of the intact liposome and associated drug into the site of disease progression. This property is differentiated from the benefits o f drug infusion technology, which are primarily concerned with maintenance of circulating blood levels of free drug. 1.6 Thesis objectives and hypotheses The aims of this thesis were to 1) characterize D S P C / C h o l and D M P C / C h o l formulations of mitoxantrone, 2) evaluate the compensating roles of drug delivery and drug release following i.v. administration of liposomal mitoxantrone, and 3) define the role of Kupffer cells and liposome mediated drug delivery to the liver in governing the efficacy of liposomal mitoxantrone used to treat liver localized cancer. Three connected hypotheses are addressed in this thesis which is focused on the development and characterization of liposomal mitoxantrone. The work emphasizes use of this formulation in the treatment of cancer that is progressing in the liver. Many groups have tried to take advantage of the natural tendency of liposomes to accumulate in the liver for the treatment of liver localized disease (Gabizon et al, 1983; Asao et al, 1992) but have met with limited success. Hepatocellular carcinoma has the highest rate of incidence 46 among all cancers worldwide. Current therapies, such as resectional therapy, radiation therapy, chemoembolization, cyrotherapy, are ineffective with remaining options being palliative for the patient. The only current course of action is focused on prevention through the use o f vaccination of the hepatitis B virus, as the incidence of hepatocellular carcinoma has been causally linked to the viral infection (Lee and K o , 1997). In addition to hepatocellular carcinoma, the liver is also a major site of metastasis. The majority of the cases are due to metastasis from colorectal carcinoma because of the gastrointestinal venous drainage to the liver. There is clearly a need to develop effective agents to treat liver cancer. It is argued that liposomal formulations should be more effective in treatment o f liver disease because these carriers accumulate in liver tissue to high levels. However, the results presented in this thesis suggest that drug delivery alone is not sufficient to treat liver localized disease. The first research chapter (Chapter 3) addresses the hypothesis that in a site where liposome accumulation is rapid, drug biological availability is more critical in defining therapeutic activity than drug delivery. Using liposome lipid composition as the primary regulator of drug release, it is demonstrated that D M P C / C h o l mitoxantrone is much more active in the treatment of liver disease in comparison to D S P C / C h o l mitoxantrone. It is concluded that mitoxantrone release is the dominating factor controlling biological activity of the liposomal drug in tissues where the rate of liposome accumulation is rapid. In Chapter 4 the question of whether drug release or liposome-mediated drug delivery becomes the dominant factor controlling therapeutic activity under conditions where the rate of liposome accumulation is slow and tumour development is within a site outside the liver is addressed. D S P C / C h o l mitoxantrone and D M P C / C h o l mitoxantrone delivery in tumours established following s.c. injection of human L S I 80 and A431 cell lines is measured and then compared to the anti-tumour activity of the drug. The results suggest that liposomal mitoxantrone induced delays in tumour growth are achieved using a 47 liposomal formulation that is selected on the basis of drug release attributes, even when the liposome accumulation rate in the site of tumour growth is slow. The research focus returns to liver localized disease in the final research chapter (Chapter 5) which documents the fact that liposomal mitoxantrone is particularly well suited for treatment of cancer that is progressing primarily in the liver. It also addresses two simple hypotheses: 1) strategies which result in reduced delivery of mitoxantrone to liver w i l l result in decreased therapeutic activity and 2) Kupffer cells play a significant role in defining the therapeutic activity of liposomal mitoxantrone. Surprisingly the second hypothesis was not supported by data that used M P S blockade to effect decreases in liver delivery of liposomal mitoxantrone. The results clearly indicate that Kupffer cells are not responsible for mediating the therapeutic activity o f liposomal mitoxantrone. In addition, the results with formulations prepared with PEG-modified lipids where the anti-tumour activity of the entrapped mitoxantrone is significantly reduced in comparison to the formulations which do not contain the lipid, imply that cell processing may be necessary for the formulation to be therapeutically active. 4 8 CHAPTER 2 MATERIALS AND METHODS 2 . 1 Materials Novan t rone® (mitoxantrone hydrochloride), Adr iamycin® (doxorubicin hydrochloride), and Oncov in® (vincristine sulphate) were obtained from the British Columbia Cancer Agency and are products of Wyeth Ayerst, (Montreal, PQ), Adr ia Laboratories (Mississauga, O N ) , and Faulding (Vaudreuil, PQ) respectively. Clodronate (dichloromethylene-bisphosphonate) was generously donated by Boehringer Manneheim (Mannheim, Germany). 1,2 Distearoyl-src-glycero-3-phosphocholine (DSPC) , 1,2 dimyristoyl-5«-glycero-3-phosphocholine ( D M P C ) , 1,2 distearoyl-^«-glycero-3-phosphatidylethanolamine-polyethylene glycol 2000 (PEG) , and egg phosphatidycholine (EPC) were purchased from Avanti Polar Lipids (Alabaster, A L ) and Northern Lipids (Vancouver, B C ) . l,r-Dioctadecyl-3,3,3',3'-tetramethylindocarbocyanine perchlorate (Di l ) was purchased from Molecular Probes (Eugene, OR) . Citric acid, 3-(4, 5-dimethylthiazol-2-yl)-2, 5-diphenyl tetrazolium bromide ( M T T ) , N-2-hydroxyethylpiperazine-N-2-ethane-sulphonic acid (HEPES) , hydrogen peroxide ( H 2 0 2 ) , Sephadex G50 (medium), nigericin and cholesterol were purchased from Sigma Chemical Company (St. Louis, M O ) . Dibasic sodium phosphate, sodium chloride, sodium citrate, and ION hydrochloric acid were obtained from Fisher Scientific (Fair Lawn, NJ) . Hank's buffer (with and without calcium and magnesium) was purchased from Stem Cel l Technologies (Vancouver, B C ) . Rat - anti mouse F4/80 antibodies and FITC conjugated goat-anti rat antibodies were purchased from Serotec (Mississauga, ON) . O .C.T . was purchased from Tissue-Tek. (Miles Inc., U S A ) . Solvable™ was obtained from N E N (New England Nuclear) Research Products (Dupont Canada, Mississauga, 49 O N ) . [ l 4 C]-Mitoxantrone, used as a tracer, was generously donated by Wyeth Ayerst (Montreal, PQ) . [ 3H]-Cholesteryl hexadecyl ether (CHE) , a lipid marker that is not exchanged or metabolized in vivo (Stein et al, 1980), and [ 3H] thymid ine were purchased from Amersham (Oakville, ON) . Aquacide II was purchased from Terochem Laboratories Ltd. (Edmonton, A B ) . A431 (a human squamous carcinoma cell line) and L S I 8 0 cells (a human colon carcinoma cell line) were purchased from the A T C C (Manassas, V A ) and maintained in culture. The L1210 and P388 tumour cell lines were originally purchased from the N C I tumour repository (Bethesda, M D ) and cells were obtained from ascites fluid generated weekly by passage in B D F 1 mice. Cells were used for experiments after the third passage and before the twentieth. Once the cells reach the twentieth passage, these care discarded and the new cell lines revert back to the original N C I tumour stock. Female C D 1 , D B A 2 and BDF1 mice (8-10 weeks old) were purchased from Charles River Laboratories (St. Constant, PC). Female S C I D / R A G - 2 were bred at the British Columbia Cancer Agency Animal Breeding Facility. 2.2 Preparation of liposomes D S P C / C h o l (55:45; mohmol), D M P C / C h o l (55:45; mokmol), and D M P C / C h o l / P E G (50:45:5;mol:mol:mol) liposomes were prepared using well established extrusion technology (Hope et al, 1985). When D i l was used as a fluorescent lipid marker, it was added at a ratio o f 0.4 mg to 100 mg total lipid. The indicated phospholipid and cholesterol mole ratios were dissolved in chloroform and dried down to a homogenous lipid film under a stream of nitrogen gas. This lipid film further was dried under vacuum for 3 hours to remove any residual chloroform. Subsequently, the lipid film was hydrated in a 300 m M citric acid buffer (pH 4.0) to a final l ipid concentration of 100 mg/ml. The resulting multilamellar vesicle mixture was frozen in liquid nitrogen and thawed five times (Mayer et al, 1986) and extruded through three 100 nm 50 (pore size) stacked polycarbonate filters (Nuclepore, Pleasanton, C A ; Poretics Corp., Mississauga, ON) using an extrusion device (Lipex Biomembranes Inc., Vancouver, B C ) . The resulting large unilamellar vesicles ( L U V s ) were sized by quasielastic light scattering using a Nicomp 270 submicron particle sizer (Pacific Scientific, Santa Barbara, C A ) operating at 632.8 nm. The mean diameter of these liposomes was 100-120 nm. 2.3 Transmembrane pH gradient loading of mitoxantrone Mitoxantrone was encapsulated using a transmembrane pH gradient driven loading procedure (Madden et al, 1990; Mayer et al, 1985). The procedure used is analogous to that employed for vincristine (Boman et al, 1993) and consisted of incubating liposomes at 65°C for 10 minutes prior to addition of sufficient mitoxantrone to achieve a final drug to lipid weight ratio of 0.1. The p H of this mixture was then increased from pH 4.0 to 7.2 by the addition of 350 ul of 0.5 M N a 2 H P 0 4 buffer to 1.0 ml of the drug liposome mixture. The resulting mixture was incubated at 65 °C for an additional 15 minutes. Encapsulation efficiency for mitoxantrone was determined at 3 different temperatures: 37 °C, 50 °C, and 65 °C using size exclusion chromatography on mini-spin columns made of Sephadex G-50 (Madden el al, 1990). Aliquots of the sample (100 ul) were taken at intervals over a 2 hour time period and assayed for drug encapsulation. Drug and lipid concentrations in the samples collected in the void volume of these columns were determined by measuring [ 3 H ] - C H E and [ i 4C]-mitoxantrone. Radioactivity was assessed by mixing the sample with 5 ml Pico-Fluor 40 (Packard, Meriden, C T ) scintillation cocktail and counted with a Packard 1900 scintillation counter (Packard, Meriden, C T ) . 51 2.4 Transmembrane p H gradient loading of vincristine Vincristine was encapsulated using a transmembrane pH gradient driven loading procedure as described by Boman et al. (1993). The procedure consisted of incubating liposomes at 65 °C for 10 minutes prior to addition of sufficient vincristine to achieve a final drug to lipid weight ratio of 0.1. The p H of this mixture was then increased from pH 4.0 to 7.2 by the addition of 350 ul of 0.5 M N a 2 H P 0 4 buffer to 1.0 ml of the drug liposome mixture. The resulting mixture was incubated at 65 °C for an additional 15 minutes. Encapsulation efficiency was approximately >95% for vincristine. 2.5 Transmembrane p H gradient loading of doxorubicin Doxorubicin ( D O X ) was encapsulated in the liposomes using the transmembrane p H gradient loading procedure (interior acidic) employing sodium carbonate as the alkalinizing agent and a drug to lipid weight ratio of 0.2:1 (Mayer et al, 1994). Empty preformed liposomes with interior p H of 4.00 (300 m M citrate buffer) were titrated with 0.5 M sodium carbonate to a p H of 7.8 -8.0. Doxorubicin, solubilized in H B S , and the titrated vesicle solution were heated at 65°C for 2 min prior to addition of doxorubicin to the liposome solution. The mixture was vortexed for 2-3 min at 65°C and then maintained at this temperature for an additional 10 min to facilitate complete drug loading. Encapsulation efficiency was > 95%. Liposomal D O X preparations were diluted with saline as necessary prior to in vivo administration. 52 2.6 Preparation of EPC/Chol clodronate liposomes Clodronate liposomes were prepared as outlined by Van Rooijen and Sanders (1994) with minor modifications. A n E P C / C h o l (86:8 wt:wt) lipid film was prepared by weighing out the required amounts o f E P C and cholesterol. Chloroform was then added to the lipids and the solution was dried down under a stream of nitrogen. The resulting film was then kept under vacuum for 3 hours. The E P C / C h o l f i lm was hydrated in a 5 ml solution of clodronate (2 mg/ml) and subsequently subjected to 5 freeze-thaw cycles in order to increase encapsulation efficiency (Mayer et al, 1986). The resulting suspension was centrifuged at 6,000 x g for 20 minutes to separate the unencapsulated clodronate from the clodronate M L V s . The M L V s form a milky band on top of the suspension. The lower suspension was removed and the liposomes were resuspended in P B S . This liposome were washed in P B S and centrifuged at 20,000 x g for 30 minutes three times. The resulting pellet of clodronate M L V s were then resuspended in 4 mis of phosphate buffered saline (PBS). 2.7 Microculture tetrazolium assay The modified microculture tetrazolium ( M T T ) assay was used to determine the IC 5o values for mitoxantrone, doxorubicin, and vincristine on L1210, LS180, and A431 cells (Al ley et al, 1988). Briefly, L1210 cells were obtained through in vivo cultivation in the mouse peritoneum. Typical ly, 10 6 cells were inoculated intraperitoneal (i.p.) and the tumour progressed for 7 days prior to isolation of cells to be used for cytotoxicity assays. Cells were isolated from the mice by peritoneal lavage and the collection of ascitic fluid into E D T A containing tubes. L1210 cells were then separated from lymphocytes and R B C s by Ficoll-Hypaque density gradient 53 centrifugation, where cells at the interface were collected and placed into R P M I 1640 medium containing 10% F B S . The cells were washed three times prior to transferring the cells into a T75 culture flask. The resulting cell suspension was incubated at 37°C in a humidified incubator with 5% CO2 for 4 hours. A l l non-adherent cells were transferred into T75 flasks and diluted to a concentration of approximately 10 5 cells/ml. The cells were incubated for 24 hours prior to use in a cytotoxicity study. A431 and L S I 8 0 cells were harvested from exponential phase cultures and counted by Trypan blue exclusion (cell preparations demonstrating viability >90% were used) prior to dispensing the cells into 96-well flat-bottomed CostarR (Cambridge, M A ) culture plates (2000 cells/100 ul /well for a 3-day incubation). The cells were exposed to defined concentrations of the anti-cancer drug (diluted with R P M I 1640 supplemented with 10% heat-inactivated F B S ) over a 3-day incubation at 37°C, 5% C 0 2 and 100% relative humidity. The M T T assay consisted of adding 50 u.1 M T T (5 mg/ml P B S , filtered through 0.45 um filter units, and stored at 4°C for not longer than 1 month) to each well and the plates were further incubated for 4 hours at 37°C. Subsequently, plates were centrifuged and the supernatant aspirated slowly through a blunt 18-gauge needle. The reaction product retained in the viable cells was thoroughly solubilized by the addition of 150 ul D M S O . The plates were read spectrophotometrically at 570 nm in a Dynex Technologies M R X multiplate reader (Dynex Technologies, Chantily, V A ) . Cytotoxicity was expressed in terms of percentage o f control absorbance (mean ± s.d.) following subtraction of background absorbance. The IC50 was determined from a plot of percentage absorbance vs. log drug concentration of the data obtained in triplicate. 54 2.8 In vitro characteristics of liposomal mitoxantrone For release studies, liposomal mitoxantrone formulations were prepared as outlined in section 3. The resulting drug loaded liposomes were transferred into 25 mm diameter Spectrapor dialysis tubing (10,000-12,000 molecular weight cut off, Spectrum Medical Industries, Los Angeles ,CA) and the samples (3 ml) were dialyzed against 1 liter of H B S at 37 °C. A t the indicated time points, 100 ul samples were taken from the dialysis bag and assayed for drug and lipid using the mini-spin columns as described above. The experiment was then repeated in the presence of nigericin, an ionophore that collapses the pH gradient by promoting exchange of a monovalent cation (eg. K + , N a + ) with H + . The ionophore was added to the sample and external buffer to a concentration of 120 n M . Further release experiments were carried out in the presence of serum. Liposomal mitoxantrone was prepared as above. Liposomal mitoxantrone (200 ul) was incubated with 800 ul of 100% fetal bovine seurm at 37°C for 24 hours. After incubation, 500 ul of the mixture was applied to a Biogel A-15 column in order to separate released drug from liposomal drug. 2.9 Plasma elimination and distribution studies Female CD1 mice (20-25 g, 4 per group) were injected with a 10 mg/kg drug dose via the lateral tail vein. A t 1, 4, 24, and 48 hours animals were terminated by C 0 2 asphyxiation and whole blood was collected via cardiac puncture and placed into E D T A coated tubes (Microtainers, Becton Dickinson). Plasma was isolated following centrifugation of whole blood at 500 x g for 10 minutes. Aliquoted plasma samples (100 ul) were mixed with 5 ml Pico-Fluor 40 and counted for [ 3H] and [ 1 4 C] . 55 Tissue weights were determined by placing (isolated and saline washed) tissues into pre-weighed glass tubes before reweighing and freezing at -70 °C. Appropriate volumes of distilled water were added to the tissues and homogenized with a Polytron tissue homogenizer (Kinematica, Switzerland) to achieve a 10% homogenate (w/v). Aliquots of the homogenate (200 u,l) were mixed with 500 u.1 of Solvable™ and incubated at 50 °C for 3 hours. After the resulting mixture was cooled to room temperature, 50 ul of 200 m M E D T A , 200 u.1 of 30% H 2 0 2 and 25 ul of 10 N HC1 were added. Five ml of Pico-Fluor 40 was added to the samples and radioactivity ( [ J H]-C D E and [ 1 4C]-mitoxantrone tracer) was determined using a Packard 1900 scintillation counter. 2.10 A431 and LS180 tumour accumulation and plasma elimination studies of liposomal mitoxantrone Female S C I D / R A G - 2 mice (18-20 g, 4 per group) were inoculated bilaterally with 2 x 106 A431 cells or 1 x 10 6 LS180 cells subcutaneously on the hind regions of the back. Once the tumours reached a measurable size (tumour volume > 0.05 cm 3 ), as measured using calipers, mice were injected with a 10 mg/kg drug dose of free mitoxantrone, D S P C / C h o l mitoxantrone, or D M P C / C h o l mitoxantrone via the lateral tail vein. A t 4, 24, 48, and 96 hours animals were terminated by C 0 2 asphyxiation and whole blood was collected via cardiac puncture and placed into E D T A coated tubes (Microtainers, Becton Dickinson). Plasma was isolated following centrifugation of whole blood at 500 x g for 10 min. Aliquoted plasma samples (100 ul) were mixed with 5 ml Pico Fluor 40 (Packard, Meriden, C T ) and [ 3H] and [ 1 4 C] were measured using a Canberra Packard 1900 scintillation counter. Isolated tissues were processed as outlined in Section 2.9. 56 2.11 Plasma elimination and biodistribution studies in MPS blockaded mice Female CD1 mice (20-25 g, 4 mice per group) were injected with a 10 mg/kg drug dose o f D M P C / C h o l mitoxantrone or D M P C / C h o l / P E G mitoxantrone via the lateral tail vein. To achieve hepatic M P S blockade to alter the plasma elimination and biodistribution of D M P C / C h o l liposomal mitoxantrone, animals were injected i.v. with a 2 mg/kg drug dose o f D S P C / C h o l doxorubicin (10 mg/kg lipid dose) 24 hours prior to injection of the D M P C / C h o l mitoxantrone. A t 1 and 4 hours, 25 u.1 of blood was collected in E D T A coated microcapillary tubes from the tail vein which had previously been given a small cut with a scalpel. Blood was mixed with 250 u.1 of 5% E D T A and spun for 15 minutes at 500 x g. The supernatant was collected and the resultant pellet was resuspended in Hanks buffered saline solution (250 ul) and spun again at 500 x g. The supernatant was collected and pooled with the first supernatant prior to addition of 5 ml of scintillation fluid. Radioactivity in the sample was assessed by scintillation counting. A t 24 hours, mice were terminated by CO2 asphyxiation, and whole blood was collected via cardiac puncture and processed as outlined in section 2.9. Livers were harvested and processed as outlined in section 2.9. 2.12 Liposome mediated drug delivery to region of tumour cell inoculation In order to measure mitoxantrone and liposomal lipid accumulations under conditions where the tumour was not established, the following protocol was used. Prior to inoculation, L S I 8 0 cells were incubated with [ 3H]-thymidine for 48 hours. The adherent cells were rinsed with R P M I media and cell suspensions were prepared by adding trypsin-EDTA followed by a brief (< 1 minute) incubation. The radiolabeled cells were then resuspended in R P M I media to a concentration of 20 x 10 6 LS180 cells/ml. Viabi l i ty was assessed using Trypan blue and cells 57 were counted using a hematocytometer. The injection sites on female S C I D / R A G - 2 mice were shaved and marked. LS180 cells (1 x 10 6 ) in 50 ul of R P M I med ia were injected bilaterally subcutaneously into the inferior dorsal region of the mice. A n equivalent amount of the cell suspension was taken for scintillation counting. Forty-eight hours after tumour cell inoculation, 10 mg/kg drug dose of D S P C / C h o l mitoxantrone, D M P C / C h o l mitoxantrone, or free mitoxantrone was injected i.v. [ 1 4C]-mitoxantrone was used as a tracer. Twenty-four hours later, mice were terminated using C 0 2 asphyxiation and blood was collected via cardiac puncture and processed as outlined in the plasma elimination studies. A 1.5 cm x 1.5 cm section of skin and underlying muscle area surrounding the inoculation site was removed and processed as outlined in section 2.9. This study was repeated using cells which were not labeled with [ 3H]-thymidine so that liposomal lipid ( [ 3 H]-CHE) and drug ([ 1 4C]-mitoxantrone) delivery to the region o f cell inoculation could be measured simultaneously. Values obtained using this technique were reported as total delivery to the site of cell inoculation. 2.13 Establishing the maximum tolerated drug dose and L1210 and P388 efficacy studies The maximum tolerated drug dose ( M T D ) was determined in limited dose ranging studies where female B D F 1 mice in groups of two were given drug by a single i.v. injection. Weight loss and signs of stress/toxicity were monitored for 30 days. If individual animals lost greater than 20% of the original body weight, they were terminated. If animals appeared severely stressed as judged by appearance and/or behavior, as assessment made by qualified animal care technicians, they were terminated. The M T D was estimated as the dose where tumour-free animals survived for a period of 30 days after drug administration. A t the end of the 30 day period, animals were terminated by C 0 2 asphyxiation and necropsies were completed to identify any additional toxicities. The exact LDio dose of the different mitoxantrone formulations was not determined as 58 such toxicity studies are not approved by the Canadian Council on Animal Care or the institutional Animal Care Committee. For L1210 and P388 efficacy studies, female B D F 1 mice (19-21 g, typically 2 sets of 5 mice per group were used providing an n value of at least 10) were injected with 104 L1210 cells or 10 5 P388 cells /.v. 24 hours before a single treatment of the indicated drug dose and formulation. When these cells are injected i.v., they seed primarily in the liver and spleen (See Chapter 3). For animals injected with L1210 cells, tumour progression is characterized by increased liver and spleen weight and histological studies indicate the presence of massive, diffuse infiltration of the liver. For animals injected with P388 cells, liver and spleen mass increase and the histopathology reveals discrete foci of tumour cells that progressively become larger over time. M i c e were given the specified drug dose in a volume of 200 ul and, where required, drug loaded liposomes were concentrated (using Aquacide II) prior to administration. The animals were monitored daily for any signs of stress and were terminated when body weight loss exceeded 20% or when the animals exhibited signs of lethargy, scruffy coat, dehydration or labored breathing. When animals were terminated, the survival time was recorded as the following day. Survival times were monitored for sixty days and drug induced increases in life span (% ILS) were calculated by dividing the median survival time of the treated by the median survival time of the control mice (saline treated). 2.14 Liposomal mitoxantrone anti-tumour efficacy using the human A431 and LS180 solid tumour models S C I D / R A G - 2 mice were inoculated bilaterally with 2 x 106 A431 cells or 2 x 10 6 LS180 cells 14 days prior to initiation of drug treatment. Tumour bearing animals (tumour size > 0.05 cm 3 ) were 59 given a single i.v. injection of free mitoxantrone, D S P C / C h o l liposomal mitoxantrone, or D M P C / C h o l liposomal mitoxantrone. Control mice were injected with saline. Previous results obtained with immunocompetent B D F 1 mice indicated that free mitoxantrone was tolerated at 10 mg/kg and liposomal formulations were tolerated at 20 mg/kg. In contrast, both liposomal drug formulations proved to be toxic (non-tumour related deaths were observed in 100% of the animals within 15 days after administration) in tumour bearing S C I D / R A G - 2 mice when administered at 20 mg/kg and free drug was toxic at the 10 mg/kg dose. It should be noted that S C I D / R A G - 2 mice were selected because they tolerated D N A damaging agents much better than other S C I D mice (e.g. Toronto SCID and N O D / S C I D mice). Based on drug dose titrations from 5 to 20 mg/kg, the maximum therapeutic dose of drug when given as a single i.v. injection was defined as 5 and 10 mg/kg for the free and liposomal drugs, respectively. Animal weights and tumour volumes were measured daily until the tumour mass exceeded 10% of the animals original body weight or until the tumours showed any sign of ulceration. Tumour volume was determined by measuring tumour dimensions and calculating volume with the equation (Tomayko and Reynolds, 1989): (TT / 6) x length x width 2 2.15 Treatment of non-established LS180 and A431 tumours In an effort to establish optimal conditions for treating S C I D / R A G - 2 mice inoculated with L S I 80 and A431 cells, studies evaluating treatment of animals two days after tumour cell inoculation were completed. Treatment was based on single (5 mg/kg free drug and 10 mg/kg liposomal drug) and multiple (1.5 mg/kg free drug and 3.5 mg/kg liposomal drug) doses. The latter consisted of intravenous injections on days 2, 3 and 4. Other dose schedules were evaluated (e.g. 60 days 2, 4 and 6; days 2, 6 and 10) but under the conditions employed, optimal therapy was obtained using days 2, 3 and 4 schedule. Control mice were injected with saline. 2.16 Efficacy of liposomal mitoxantrone in the Lv. L1210 tumour model with and without MPS blockade Female B D F 1 or D B A 2 mice were inoculated with 1 x 104 L1210 tumour cells i.v. and 24 hours after tumour cell inoculation mice were treated with a 10 mg/kg drug dose o f D M P C / C h o l mitoxantrone or D M P C / C h o l / P E G mitoxantrone. M i c e were given the specified drug dose in a volume of 200 ul . In order to assess the impact of hepatic M P S blockade therapeutic activity, mice were injected i.v. with either DSPC/Cho l doxorubicin (2mg/kg drug), D S P C / C h o l vincristine ( lmg/kg drug), or E P C / C h o l clodronate 2 hours after tumour cell inoculation. Controls indicated that the agents used to blockade the hepatic M P S blockade had no therapeutic activity at the doses administered. The animals were monitored and terminated as described in Section 2.13. 2.17 F4/80 staining of macrophages in the liver CD1 mice were injected with either 2 mg/kg drug dose of D S P C / C h o l doxorubicin or E P C / C h o l clodronate. Control livers were left untreated. Twenty-four hours after treatment, mice were terminated via C 0 2 asphyxiation and livers were harvested. The livers were rinsed in ice cold P B S buffer and placed in O.C.T. embedding solution for 30 minutes before frozen at -70 °C. Cryostat sections (5 um)were prepared using a Frigocut 2800E microtome from Leica. The slides were then washed in P B S and incubated with the rat-anti mouse F4/80 primary antibody and then a F I T C goat anti-rat secondary antibody. A Leitz Dialux fluorescence microscope (at 61 40 x magnification) was used to evaluate F ITC fluorescence of the sections (430-490 nm cut off filter) with fluorescent photomicrographs obtained using a Orthomat microscope camera. A l l images were recorded on Fuji color A S A 4 0 0 negative film. 2.18 Hepatocyte isolation Hepatocytes were isolated from female CD1 mice as described by (Klaunig et al, 1981) with slight modification. M i c e were terminated via C 0 2 asphyxiation. Livers were harvested and kept in ice cold Hank's buffer (without calcium and magnesium). Using two scalpel blades, the livers were minced to a fine mixture and this was transferred to a 15 ml culture tube. Hank's buffer (without calcium and magnesium) was added to final volume of 5 ml. Three hundred ul of collagenase (4 mg of collagenase/ml of Hank's with calcium and magnesium) was then added to the solution and incubated on a rotating tube rack at 37°C for 30 minutes. The resulting solution was then strained through a 40 um nylon filter and 40 ml of Hank's buffer added. This was spun for 1 minute at 50 x g. The supernatant was extracted and the resulting pellet was reconstituted in another 40 ml of Hank's buffer (without calcium or magnesium). This solution was spun for 1 minute at 50 x g. This step was repeated twice. The final pellet was reconstituted in 5 ml of Hank's buffer (without calcium and magnesium). Viabi l i ty was assessed using Trypan Blue and was found to be greater than 90%. Hepatocytes were counted using a Coulter cell counter Z M 901 (Coulter, Burlington, ON) . 2.19 Confocal microscopy D i l [a fluorescent lipid label that is not exchanged or metabolized (Claassen, 1992; Honig and Hume, 1986)] was added to liposomes as described in Section 2.2. D M P C / Choi , 62 D M P C / C h o l / P E G liposomes were loaded with mitoxantrone and injected at a drug dose of 10 mg/kg (100 mg/kg lipid dose). Twenty-four hours after injection, mice were terminated and the livers were gentlely harvested and rinsed in P B S . Subsequently, the livers were placed in O.C.T . for 30 minutes before freezing at - 7 0 ° C . Sections (5 um in thickness) were made using a cryostat and imaged using confocal micoscopy. Confocal images were collected on a Optiphot 2 research microscope (Nikon Japan) attached to a confocal laser scanning microscope (MRC-600 , BioRad Laboratories, Hercules C A ) using C O M O S software (BioRad Laboratories). The laser line on the krypton/argon laser was 488 nm. Filterblock B H S (568 nm) was used to detect D i l (549 nm excitation, 565 nm emission). The numerical aperture was 0.75 on the lOx air objective and 1.2 on the 60x oil objective. The images were captured such that the xyz dimensions were 0.4 mm cubed (20x) and 0.2 mm pixel (60x). N I H Image version 1.61 was used for image analysis, and all images were based on maximum intensity projection. Projections made in N I H image were saved in TIFF format and then imported to Adobe Photoshop version 4.0 where final modifications were performed. 2.20 Statistical analysis A N O V A (analysis of variance) was performed on the results obtained after administration o f the two liposomal formulations and free mitoxantrone. Common time points were compared using the post hoc comparison of means, Scheffe test. Differences were considered significant at p < 0.05. Area under the curve analysis was performed using trapezoidal integration from the time points indicated. The zero point is a theorical point and was calculated as the injected dose over the plasma compartment of the mouse and then corrected for 100 u.1. 63 CHAPTER 3 INFLUENCE OF DRUG RELEASE CHARACTERISTICS ON THE THERAPEUTIC ACTIVITY OF LIPOSOMAL MITOXANTRONE 3.1 Introduction: It is wel l established that the therapeutic activity of anti-cancer agents can be improved through application of liposomal drug carrier technology (Fielding, 1991; Sugarman and Perez-Soler, 1992; K i m , 1993). In general, liposomes engender pharmacokinetic and biodistribution characteristics which lead to increases in therapeutic activity and/or reductions in drug related toxicities (Fielding, 1991; Mayer et al, 1994). Although the mechanism of therapeutic activity for liposomal anti-cancer drugs is not well understood, studies have suggested that increased drug exposure at the site of tumour growth is important (Gabizon and Papahadjopoulos, 1988; W u et al, 1993; Ba l ly et al, 1994; Mayer et al, 1994). These increases in tumour drug levels result from preferential accumulation of the liposome carrier within tumours (Gabizon, 1992; W u et al, 1993; Ba l ly et al, 1994; Ogihara-Umeda et al, 1994; Uchiyama et al, 1995). It is important to note, however, that there is no evidence suggesting that the encapsulated form of the drug is therapeutically active. It is postulated, therefore, that anti-tumour activity is mediated by drug released from regionally localized liposomes (Mayer et al, 1994). The emphasis of investigators developing liposomal anti-cancer agents has been, for the reasons cited above, on the use of liposomal lipid compositions that are less permeable to the encapsulated agent and exhibit increased circulation lifetimes. Liposomes that are retained in the plasma compartment for extended periods of time exhibit a greater tendency to accumulate in regions o f tumour growth (Gabizon and Papahadjopoulous, 1988; Gabizon, 1992; W u et al, 64 1993). However, the kinetics of this extravasation process, where liposomes leave the blood compartment and enter an extravascular site, are slow (Nagy et al, 1989; Bal ly et al, 1994). Efficient drug delivery can, therefore, only be achieved with liposomes that effectively retain the drug following systemic administration. The problem that arises through applications of liposomal carriers that are optimized for enhanced drug retention concerns evidence from studies with liposomal doxorubicin that demonstrate reduced therapeutic activity, despite efficient delivery o f drug to tumours (Parr et al.,\991). A balance between doxorubicin retention (to maximize drug accumulation in a site of tumour growth) and release (to effect therapy) has not been established. Attempts to improve the therapeutic properties of liposomal doxorubicin formulations through changes in drug release characteristics have been unsuccessful due to specific adverse effects of free doxorubicin, including cardiotoxicity (Minow et al, 1975) and drug mediated free radical damage (Rajagopalan et al, 1988). More specifically, effective modulation of doxorubicin release rates has been achieved with relatively simple changes in liposomal lipid composition (Mayer et al, 1989; Bal ly, et al, 1990); however, liposomal formulations of doxorubicin that release drug following i.v. administration, exhibit enhanced toxicity and increased doxorubicin accumulation in cardiac tissue. This effect is most dramatic for doxorubicin formulations prepared using D M P C / C h o l liposomes, which release greater than 90% of the encapsulated contents in the blood compartment within 24 hours after i.v. administration and are 3 times more toxic than free drug (Mayer et al, 1994). The studies in this chapter examine the influence of liposome drug release properties on the biological activity of mitoxantrone. The rationale for selecting mitoxantrone is based on the fact that this drug is less cardiotoxic than doxorubicin (Weiss, 1989) and is not capable o f generating 65 free radical damage in non-proliferating cells (Durr, 1984). It is demonstrated that the in vivo rate of mitoxantrone release from D M P C / C h o l liposomes is at least 68-fold greater than that obtained from D S P C / C h o l liposomes. The pharmacodynamic characteristics of these formulations have been characterized using murine tumour models where the primary site of tumour progression is in the liver. The data illustrate how a balance between drug release characteristics and liposome mediated drug delivery to sites of tumour progression is required for optimal therapeutic activity. 3.2 Results 3.2.1 In vitro mitoxantrone uptake and release characteristics Studies evaluating in vitro drug accumulation in liposomes prepared from D M P C (Cu) /Choi and D S P C (Cis) /Choi at 37 °C, 50 °C and 65 °C are shown in Figure 3.1. A t 37 °C, less than 15% of the drug was encapsulated in either liposome formulation over the 2 hour time course. In contrast, >98% of the drug was efficiently entrapped when the incubation temperature was increased to above 50 °C. The time required to achieve maximum uptake was 45 minutes and less than 5 minutes when the incubation temperature was 50 °C and 65 °C, respectively. Uptake rate was enhanced slightly at 50 °C for the D M P C / C h o l when compared to the D S P C / C h o l systems. The results suggest that the phase transition temperature (T c ) of the phospholipid species does not markedly affect mitoxantrone loading characteristics. This result is consistent with the in vitro drug release studies (Figure 3.2) that demonstrate no difference in drug release from either liposomal formulation. The in vitro release assay used is based on dialysis against a large volume (1 L ) of H B S buffer. Under these conditions, free mitoxantrone equilibrates across the dialysis membrane in less than 8 hours. In contrast, less than 2% drug release was observed 66 Figure 3.1 Effect of temperature on pH gradient loading of mitoxantrone into DSPC/Chol (A) and DMPC/Chol (B) liposomes Loading was evaluated at three temperatures: 37 °C ( • ) ; 50 °C ( A ) ; 65 °C ( T ) . A t time zero, mitoxantrone and the liposomes were mixed together at a drug to lipid ratio o f 0.1 (wt:wt). Encapsulated drug was determined by the mini spin column procedure described in Chapter 2, section 8. Duplicate samples were taken and [ 3 H ] - C D E and [ 1 4C]-mitoxantrone were measured. Data represents the average values + S.D. of four measurements. 0 15 30 45 60 75 90 105 120 Time (minutes) 67 from the liposomal formulations over a 72 hour incubation period at 37 °C. Figure 3.2 also incorporates data obtained for mitoxantrone loaded liposomes incubated with nigericin, a H +/monovalent cation exchanger (dashed lines). Although drug release rates were increased in the presence of nigericin, there were minor differences in release rates observed for the two liposomal systems studied. After the 48 hour incubation period the D M P C / C h o l liposomes released less than 30% of the encapsulated drug in comparison to 20% drug release observed for the D S P C / C h o l system. Figure 3.3 demonstrates release of mitoxantrone from D M P C / C h o l liposomes after incubation with fetal bovine serum for twenty-four hours. The results in figure 3.3B demonstrate that the presence of serum proteins also did not enhance mitoxantrone release from D M P C / C h o l liposomes. 3.2.2 In vivo plasma elimination of liposomal lipid and mitoxantrone Results in Figure 3.4 show that the plasma elimination of liposomal lipid, following i.v. administration of mitoxantrone loaded D M P C / C h o l and DSPC/Cho l liposomes, is similar (Figure 3.4A). A n estimation of the amount of mitoxantrone retained in the liposomes remaining in the circulation can be made by determining the ratio of mitoxantrone to lipid at the indicated time points; an estimation that assumes the level of free drug in the plasma of animals given liposomal mitoxantrone is negligible. The results shown in Figure 3.4B demonstrated greater release of mitoxantrone from D M P C / C h o l liposomes than D S P C / C h o l liposomes (p < 0.05 for 24 and 48 hour time points). For D M P C / C h o l liposomes, 73% of the mitoxantrone originally associated with the carrier was released within 48 hours. In contrast, less than 5% of the drug was released from D S P C / C h o l liposomes. Between the 4 and 48 hour time points, the rate o f mitoxantrone release was estimated to be 1.7 and < 0.025 ug lipid/1 OOu.1 plasma/hour for D M P C / C h o l and D S P C / C h o l liposomes, respectively. These results are consistent with those obtained using 68 Figure 3.2 Release of mitoxantrone from DSPC/Chol (•)and DMPC/Chol (•) liposomes in HEPES buffered saline at 37 °C Solid lines indicate the absence of Nigericin. Dashed lines indicate the addition of Nigericin at time zero. Samples (100 ul) were taken from the dialysis bags and applied to Sephadex G-50 mini spin columns in duplicate and spun at 500 xg for 2 minutes. Duplicate samples were taken from the resulting mixture and [ 3H] and [ l 4 C] were measured as described in Chapter 2, section 8. Data represents the average values + SD of at least four measurements for studies in the presence of Niger ic in . 69 Figure 3.3 Release of mitoxantrone from DMPC/Chol liposomes incubated with fetal bovine serum at 37°C for 24 hours D M P C / C h o l liposomes (200 ul) were incubated with 800 ul of fetal bovine serum for 24 hours at 37°C. The solution (500ul) was applied to a B ioGe l A-15 column and fractions were collected. Panel A represents the fractions collected when empty D M P C / C h o l liposomes ( • ) and free mitoxantrone [mitox(A)] were passed down the column. Panel B represents the lipid ( • ) and mitoxantrone (A) fractions collected with loaded liposomes passed down the column. 70 entrapped doxorubicin (Mayer et al, 1994) and clearly demonstrate that control of in vivo mitoxantrone release rates can be achieved through simple changes in liposomal lipid composition. It should be noted that plasma drug levels obtained following administration o f free drug are significantly less than those obtained with the liposomal formulations. This is indicated in Figure 3.4C, a plot of plasma drug levels measured following i.v. administration of the indicated formulation. Trapezoidal area-under-the-curve ( A U C ) analysis of these plasma drug levels, from 1 to 48 hours, indicate plasma A U C s of 0.01, 167.86 and 229.86 ug mitox/lOOu.1 plasma/hour following administration of free mitoxantrone, D M P C / C h o l mitoxantrone and D S P C / C h o l mitoxantrone, respectively. 3.2.3 Acute toxicity of free and liposomal mitoxantrone Formal LDio and L D 5 0 studies are not sanctioned by the Canadian Council of Animal Care; therefore, toxic dose range finding studies in tumour free female B D F 1 mice were conducted using only 2 mice per dose. These limited dose escalation studies suggested that the M T D of free drug was approximately 10 mg/kg. When drug was encapsulated in D S P C / C h o l or D M P C / C h o l , the M T D increased to approximately 30 mg/kg. At this dose, 100% of the animals treated survived for greater than 30 days. Necropsies suggested no gross abnormalities in any of the tissues examined. A n evaluation of drug induced weight loss, however, suggested that the D M P C / C h o l liposomal formulation was more toxic than the D S P C / C h o l system. This result was confirmed in efficacy experiments, where changes in weight were measured over 14 days following initiation of treatment. For animals given 10 4 L1210 cells /.v. and treated 24 hours later with mitoxantrone, the maximum therapeutic dose of free and liposomal mitoxantrone was 10 and 20 mg/kg, respectively. The nadir in weight loss following treatment of tumour bearing animals occurred between day 12 and 13 and at this time point animals treated with free drug (10 71 Figure 3.4 In vivo release of mitoxantrone from DSPC/Chol ( • ) , DMPC/Chol ( • ) liposomes, and free mitoxantrone (A) Liposomes were loaded with mitoxantrone at a drug to lipid weight ratio of 0.1 (wt:wt). Female CD1 mice were injected at a 10 mg/kg drug dose i.v. via lateral tail vein. Panel A shows elimination of lipid from the plasma compartment over 48 hours. Panel B shows the change in the drug to lipid ratio over the 48 hour time period. Panel C shows the elimination of the free drug from the plasma compartment over 48 hours. Data represents the mean and S . E . M . obtained from at least 4 animals. (*) indicates p < 0.05. 72 73 mg/kg) lost almost 30% of their original body weight and had to be killed. In contrast, animals treated with D S P C / C h o l and D M P C / C h o l mitoxantrone (20 mg/kg) exhibited a body weight loss of 8%> and 25%, respectively. 3.2.4 L1210 and P388 anti-tumour activity of free and liposomal mitoxantrone The murine tumour models used for evaluating the anti-tumour activity of liposomal mitoxantrone were based on i.v. injection of L1210 or P388 cells. Although these cells are typically used to initiate ascitic tumours following i.p. inoculation, the cells can also be given by alternate routes of administration. When given i.v., primary sites of cell seeding include the liver and spleen. Evidence to support this is provided in Figures 3.5 and 3.6. Seven days following i.v. inoculation of 104 L1210 cells, the liver and spleen of the recipient animals showed greater than a 2- and 3-fold increase in liver and spleen weight, respectively as shown in Figure 3.5. Untreated animals must be terminated as a result of significant tumour related disease within 10 days. Histological studies indicated the presence of massive, diffuse cell infiltration throughout the liver. There were no other gross abnormalities in any other organs or tissues derived from these animals. For mice injected with P388 cells, liver and spleen weight increases were also observed. The histopathology, however, revealed discrete foci of tumour cells that progressively became larger over a 7 day time course (Figure 3.6). These i.v. tumour models were typically not responsive to chemotherapy with doxorubicin or vincristine [free or liposomally encapsulated drug (Refer to Chapter 5, Table 5.1)], hence these models were employed as a stringent measure of mitoxantrone anti-tumour activity. 74 Figure 3.5 L i v e r and spleen weights of untreated B D F 1 mice (open bars) and B D F mice previously (7 days) injected i.v. with 10 4 L1210 cells (hatched bars) On day 7, livers and spleens were taken from B D F 1 mice and weighed. The results were obtained from 4 animals and error bars indicate the S . E . M . 75 Figure 3.6 Progression of the P388 / .v . tumour model in the liver Hematoxylin and eosin staining of paraffin embedded livers of untreated B D F 1 mice (Panel A ) , and B D F 1 mice previously injected i.v. 1 day (Panel B) , 3 day (Panel C) , and 7 days (Panel D) with 10 5 P388 cells. Structural features are pointed out as: H - Hepatocytes, S- Sinusoid, V -Blood vessel, rbc - red blood cell , K - Kupffer cell. Arrowheads indicate inflammatory infiltrate and arrows show the disorganization and lack of hepatocytes during tumour cell invasion. The bar in Panel C represents 30 um. 7 6 The L1210 anti-tumour studies summarized in Table 3.1 and Figure 3.7A clearly demonstrate that the D M P C / C h o l liposomal formulation was therapeutically more active than free drug and drug encapsulated in D S P C / C h o l liposomes. As shown.in Table 3.1, the maximum % ILS achieved with free drug was 98%. Enhanced therapy was observed for drug encapsulated in D S P C / C h o l liposomes, where a maximum % ILS value of 189 was obtained at a dose o f 20 mg/kg. Improved therapy achieved with D S P C / C h o l liposomal drug was primarily a consequence of liposome mediated reductions in drug toxicity. A t 10 mg/kg, for example, the L1210 anti-tumour activity of this liposomal formulation was significantly lower than that obtained with free drug. Remarkably, treatment with D M P C / C h o l liposomal mitoxantrone resulted in 100% long term (>60 day) survival at drug doses of 10 and 20 mg/kg. The survival curves obtained for animals treated at a dose of 10 mg/kg (Figure 3.7A) clearly show that the therapeutic activity of mitoxantrone was significantly enhanced when encapsulated in D M P C / C h o l liposomes. These results were confirmed using a similar tumour model derived following z'.v. injection of P388 cells. These results, shown in Figure 3.7B, demonstrate that animals treated with the D M P C / C h o l liposomal mitoxantrone formulation were effectively cured when the drug was administered at a dose of 10 mg/kg. 3.2.4 Drug and liposomal lipid uptake in liver The results presented to this point demonstrate that 1) the rate of mitoxantrone release from D M P C / C h o l liposomes following i.v. administration was significantly greater than that measured for D S P C / C h o l liposomes and 2) D M P C / C h o l liposomal mitoxantrone was significantly more efficacious than free drug or D S P C / C h o l liposomal mitoxantrone when tested against a tumour model where the primary site of disease progression is in the liver and spleen. It has been proposed that differences in the therapeutic activity of encapsulated anti-cancer drugs w i l l be a 77 Table 3.1: L1210 anti-tumour activity of free and liposomal mitoxantrone in BDF1 mice Sample Drug Dose Lipid Dose 60 Day Mean Survival %ILS a L/Fc (mg/kg) (mg/kg) Survival (days) Control - - 0/25 8.7 N / A N / A Free Mitoxantrone 5 - 0/10 13.6 56 N / A 10 - 0/10 17.2 98 N / A 20 - 0/5 12.6 45 N / A D S P C / C h o l 10 100 0/9 14.7 69 0.85 20 200 0/10 25.1 189 1.99 D M P C / C h o l 5 50 0/5 17.2 98 1.26 10 100 10/10 >60 N D b N D 20 200 10/10 >60 N D N D a Percentage ILS (Increase in Life Span) Values were determined from mean survival times of treated and untreated control groups. If the animals survived more than 60 days the % ILS was not determined b N D can not be determined based on a 100% survival rate for 60 days c L/F (Liposomal/Free) values were calculated by dividing the mean survival time of the liposomal formulation by the mean survival time of the free drug at the equivalent dose. consequence of liposomal characteristics that regulate the drug exposure within sites of disease progression. Therefore, in addition to assessing drug release from liposomes in the plasma compartment, it is also important to correlate anti-tumour activity with drug levels at the site of tumour progression. For this reason, drug delivery to the liver, a primary site of disease progression for the i.v. tumour models employed, was evaluated. Results, shown in Figure 3.7, were obtained in tumour free CD1 mice. It should be noted that drug/liposome plasma elimination and biodistribution data were similar in tumour free CD1 and tumour bearing B D F 1 mice. A s shown in Figure 3.8A, liposomal lipid accumulation in the liver was similar for both D S P C / C h o l and D M P C / C h o l liposomal mitoxantrone formulations over 48 hours. Unl ike doxorubicin (Bal ly et al, 1990), the presence of entrapped mitoxantrone did not cause significant 78 reductions in liposomal lipid accumulation in the liver. Empty D M P C / C h o l liposomal lipid uptake in the liver, for example, was not significantly different from mitoxantrone loaded D M P C / C h o l liposomes. Figure 3.8B demonstrates that the level of mitoxantrone achieved in the liver following i.v. administration of D M P C / C h o l liposomal mitoxantrone is less than that observed for D S P C / C h o l mitoxantrone (p < 0.01 for the 48 hour time point). A U C analysis of liver drug levels, from 1 to 48 hours, indicates liver A U C s of 2564, 1810, and 1070 ug drug/g liver/hr following i.v. administration of D S P C / C h o l mitoxantrone, D M P C / C h o l mitoxantrone, and free mitoxantrone, respectively. Notably the liposomal formulation that engenders the greatest level of drug exposure in the liver (DSPC/Chol ) did not provide the greatest therapeutic benefit. 3.3 Discussion The therapeutic index of most anti-cancer drugs is narrow, with severe toxic side effects occurring within the same dose range required to mediate effective therapy. Although a variety of experimental strategies have been developed to improve the therapeutic index of anti-cancer drugs, these strategies have a common aim: to improve drug specificity. The principle benefit postulated for the use of liposomes as carriers of anti-cancer drugs is liposome mediated increases in drug delivery to the disease site and decreases in drug delivery to healthy tissues and organs (Sugarman and Perez-Soler, 1992; Mayer et al, 1994). Using this as a rationale, emphasis is placed on the importance of designing liposomes that have a greater propensity to accumulate within disease sites (Gabizon and Papahadjopoulos, 1988; Gabizon, 1992; Ogihara-Umeda et al, 1994; Uchiyama et al, 1995). In this regard, liposome carriers have been optimized with respect to maximizing the amount of drug contained per liposome (Mayer et al, 79 Figure 3.7 Survival times of BDF1 mice injected with 104 L1210 cells (Panel A) or 10s P388 cells (Panel B) iv. via the lateral tail vein and treated with mitoxantrone Twenty four hours after tumour cell inoculation, the mice were treated with 10 mg/kg dose of free mitoxantrone ( A ) , D S P C / C h o l ( • ) , and D M P C / C h o l ( • ) liposomal formulations. Untreated (saline) animals served as controls ( T ) . 100 80 A 60 CD CO o o <D C O CO -t—• c CD o 0) CL 40 20 100 80 60 B 40 A 20 A -rt 1 — . 10 20 30 40 50 Days after Tumor Cell Inoculation 60 80 Figure 3.8 Lipid and drug levels in the liver of mice after injection of DSPC/Chol mitoxantrone (•), DMPC/Chol mitoxantrone(B), empty DMPC/Chol liposomes (•), and free mitoxantrone (A) The liposomal lipid dose was 100 mg/kg and the drug dose was 10 mg/kg. Panel A shows the amount o f lipid per gram of liver and panel B shows the amount of mitoxantrone per gram of liver measured over 48 hours. Drug and lipid levels were determined as described in Chapter 2, section 9. The data represents the mean + S . E . M . from at least 3 animals. (**) indicate p < 0.01 when compared to free drug. 1000 0 8 16 24 32 40 48 Time (Hours) 81 1989; Mayer et al., 1994), increasing drug retention characteristics (Mayer et al, 1989; Boman et al, 1994) and augmenting the circulation lifetime of the drug loaded carrier (Gabizon, 1992; W u et al, 1993). However, it can be suggested that the therapeutically active component of a liposomal anti-cancer drug formulation is the free drug. It is believed that the primary source of free drug arises from regionally localized liposomes (Mayer et al, 1994). Therefore, this research has attempted to establish a balance between efficient liposome delivery to the disease site and controlled drug release. The latter can be achieved for certain drugs by changing the liposomal lipid composition (Boman et al, 1994; Mayer et al, 1994;). This study illustrates how controlled drug release can engender significant improvements in therapeutic activity of the anti-cancer drug mitoxantrone. Tt was surprising that differences in drug accumulation and leakage rates for D S P C / C h o l and D M P C / C h o l liposomes were not substantial when evaluated in vitro, even when the liposomes were incubated in the presence of nigericin. The phase transition temperatures (T c ) for D S P C and D M P C are 55.3 °C and 23.9 °C, respectively (Lewis el al, 1987) and it was anticipated that differences in the gel to liquid crystalline phase transition of these phospholipids would be reflected by changes in permeability characteristics. This was evident for liposomal formulations of vincristine, where a good correlation between phospholipid T c and drug leakage, in vitro, was observed (Boman et al, 1993). Collapse of the transmembrane p H gradient did increase drug release from the liposomal formulations; however, no substantial differences in the rate of drug release from the D S P C / C h o l and D M P C / C h o l liposomes were noted. Fol lowing pH gradient mediated uptake, it is believed that drugs such as mitoxantrone can form insoluble precipitates within the liposome (Madden et al, 1990). If this is the case, permeability characteristics o f the drug in a precipitated form may be less dependent on membrane characteristics or the presence of 82 a residual transmembrane p H gradient. It is not understood, however, why differences in drug permeability become apparent in vivo. Mitoxantrone was selected as a model drug for these studies for two reasons. First, the drug loading and release characteristics of mitoxantrone are comparable to doxorubicin (Madden et al, 1990). Second, mitoxantrone is less cardiotoxic than doxorubicin (Dukart et al, 1985; Neidhart et al, 1986; Bennett et al, 1988; Weiss et al, 1989). Liposome mediated increases in mitoxantrone M T D observed in this report are comparable to those reported for a liposomal mitoxantrone formulation prepared using an anionic lipid-drug complex (Schwendener et al, 1991; Schwendener et al, 1994). The liposomal formulations evaluated here, however, exhibit significantly better drug retention characteristics than those formulations described by Schwendener et al. This is reflected in higher blood levels and improved circulation lifetimes for mitoxantrone encapsulated in the PC/Cho i based liposomal carriers. Differences in drug release characteristics may be a consequence of the use of anionic lipids. Anionic lipids w i l l increase liposome elimination rates (Hwang, 1987) and have been shown to enhance release of the anthracycline doxorubicin even when encapsulated using the transmembrane p H gradient loading procedure (Mayer et al, 1989). Clearly, when the rate of drug dissociation from the liposomal carrier is very rapid, carrier mediated changes in drug pharmacokinetics and biodistribution w i l l not be significant and changes in biological activity (relative to drug administered in free form) w i l l be minimal. Studies evaluating the therapeutic activity of D S P C / C h o l and D M P C / C h o l liposomal mitoxantrone (Figure 3.7 and Table 3.1) establish that both drug delivery and drug release are important attributes of an optimal liposomal anti-cancer drug formulation. The i.v. L1210 tumour model was selected for these studies, in part, because L1210 cells seed primarily in the 83 liver and spleen following i.v. administration. It is well established that these tissues are primary sites of liposome accumulation (Hwang, 1987; Sugarman and Perez-Soler, 1992). Further, other investigators have shown using experimental models of liver cancer that the therapeutic activity of liposomal formulations of a novel platinum compound and doxorubicin analogue is enhanced compared to free drug (Perez-Soler, 1989; Gabizon, 1992). It is perplexing, therefore, that models o f liver cancer have not been used more frequently to characterize the pharmacodynamic behavior of liposomal anti-cancer drugs. These studies have shown that mitoxantrone delivery to the liver is enhanced when using D S P C / C h o l liposomes in comparison to D M P C / C h o l liposomes (see Figure 3.8B). Increased liposomal drug exposure in this tissue, however, does not result in improved therapeutic activity. In fact, the D M P C / C h o l liposomal formulation, which exhibits controlled release characteristics and a reduced capacity to deliver drug to the liver, was significantly more effective. Thus, it is not sufficient to develop drug carriers that accumulate at the disease site in high levels, one must also engineer appropriate drug release rates. Studies completed and summarized in Chapter 5 have demonstrated using the i.v. L1210 tumour model that E P C / C h o l liposomal doxorubicin, D S P C / C h o l liposomal doxorubicin, and D S P C / C h o l liposomal vincristine are relatively ineffective in treating this model, typically producing increases in lifespan of less than 50% at the maximum therapeutic doses (see Table 5.1). A possible explanation for the effectiveness of liposomal mitoxantrone may be related to the fact that this encapsulated drug does not appear to affect the liver Kupffer cells. These studies have shown, for example, that empty and mitoxantrone loaded liposomes exhibit comparable plasma elimination profiles and comparable levels of uptake in liver (see Figure 3.8). This is contrary to effects observed with vincristine (Boman et al., 1994) or doxorubicin (Bal ly et al, 1990) loaded liposomes, where encapsulated drug significantly increases the circulation lifetime of the liposomal carrier. This effect is due, in part, to drug mediated blockade of 84 phagocytic cells in the liver. It can be suggested that the blockade effect may adversely affect the therapeutic activity of liposomal anti-cancer drugs in treating tumours that are progressing in the liver and that phagocytic cells in the liver may have a significant role in defining the anti-tumour activity o f liposomal mitoxantrone. In conclusion, a liposomal mitoxantrone formulation has been developed which has significant therapeutic activity. The plasma elimination curves and biodistribution data demonstrate that effective control of both drug release characteristics and target site delivery can work synergistically to achieve optimal therapy. The research described in the following two chapters w i l l continue to study liposomal formulations of mitoxantrone with the aims of: 1) further improving the therapeutic index of the drug; 2) targeting the liposomal drug for use in treatment of specific cancers, such as hepatocellular carcinomas and/or 3) developing novel formulations that effect delivery of the drug loaded carrier to tumour cells, thereby, bypassing normal cellular drug uptake mechanisms. The D M P C / C h o l formulation described here meets the first objective. 85 CHAPTER 4 FACTORS AFFECTING THE THERAPEUTIC ACTIVITY OF LIPOSOMAL MITOXANTRONE FOLLOWING INTRAVENOUS ADMINISTRATION IN SCID/RAG2 MICE BEARING ESTABLISHED HUMAN A431 AND LSI80 SOLID TUMOURS: DRUG RELEASE VERSUS LIPOSOME MEDIATED DRUG DELIVERY 4.1 Introduction Liposome formulations developed in an effort to enhance the therapeutic properties of anti-cancer drugs have traditionally focused on lipid compositions that allow for retention o f the liposomes in the circulation for extended periods of time and exhibit slow drug release rates (Mayer et al, 1989; Mayer et al, 1993; Boman et al, 1994; Gabizon et al, 1996). This strategy has been pursued based on a putative biological mechanism relying on the inherent ability of liposomes to be preferentially taken up in disease sites such as tumours (Proffitt et al, 1983; Mayer et al, 1990; Al l en et al, 1991; Bal ly et al, 1994). This uptake is interrelated with increases in tumour blood vessel permeability that occur as a consequence of angiogenesis and associated expression of vascular endothelial growth factor/vascular permeability factor ( V E G F / V P F ) in tumours (Folkman, 1985; Dvorak et al, 1988; Dvorak et al, 1991). Given the emphasis placed on maximizing liposome-mediated drug delivery to tumours, this chapter assesses the role of liposome delivery compared to drug release from liposomes in enhancing the therapeutic activity of associated anti-cancer drugs. This research has focused on the anti-cancer drug mitoxantrone for several reasons including data that suggests that mitoxantrone is: 1) less cardiotoxic compared to doxorubicin (Dukart et al, 1985; Neidhart et al, 1986; Bennett et al, 1988; Weiss et al, 1989) and 2) effective in the treatment of breast cancer, leukemia, and lymphoma (Smith et al, 1983; Durr, R . B . , 1984). Mitoxantrone has proved to be a suitable substitute for doxorubicin in clinical settings where alopecia and/or 86 cardiotoxicity are concerns (Dukart et al, 1985; Bennett et al, 1988; Neidhart et al, 1986; Weiss et al, 1989). Another property of mitoxantrone which makes it an ideal choice for the pharmacodynamic studies developed in this thesis is that the encapsulated drug does not influence the plasma elimination and biodistribution characteristics of the liposomal carrier (See Chapter 3). This is in contrast to other anti-cancer drugs such as vincristine and doxorubicin, which, when encapsulated in liposomes, engender reductions in elimination rate of the associated carrier following intravenous administration. This effect has been attributed to a direct toxicity of the encapsulated drug on phagocytic cells that play an important role in effecting liposome elimination from the plasma (Bally et al, 1990; Boman et al, 1994; Daemen et al, 1995). For this reason, mitoxantrone biodistribution and elimination parameters are dictated solely by attributes of the liposomal carrier rather than by combined effects induced by encapsulated-drug dependent changes in liposome pharmacokinetic behavior. Using liposomal formulations of mitoxantrone differing in their in vivo drug retention characteristics, it was demonstrated in Chapter 3 that drug release is required for optimal therapeutic activity when the tumour model grows in the liver. The previous chapter addressed a hypothesis suggesting that drug release is the dominating factor controlling biological activity o f liposomal drugs in tissues where the rate of liposome accumulation is rapid. The studies in this chapter addresses the question of whether drug release or liposome-mediated drug delivery becomes the dominant factor controlling therapeutic activity under conditions where the rate of liposome accumulation is slow and tumour development is within a site outside the liver. D S P C / C h o l mitoxantrone and D M P C / C h o l mitoxantrone delivery were evaluated in tumours established following s.c. injection of human L S I 80 and A431 cell lines. These cell lines were 87 selected on the basis o f empirical observations that indicated more rapid liposome uptake in L S I 8 0 tumours compared to A431 tumours. The results suggest that delays in tumour growth induced by liposomal mitoxantrone are achieved using a liposomal formulation that is selected on the basis of drug release attributes, even when the liposome accumulation rate in the site of tumour growth is slow. 4.2 Results 4.2.1 L i p i d and drug accumulation in solid L S I 8 0 and A431 tumours. L i p i d and drug levels were measured in established (> 0.05 cm 3 ) A431 and L S I 8 0 solid tumours over a 96 hour time period following a single i.v. injection of free mitoxantrone (10 mg/kg), D S P C / C h o l liposomal mitoxantrone (10 mg drug/kg, 100 mg total lipid/kg) and D M P C / C h o l liposomal mitoxantrone (10 mg drug/kg, 100 mg total lipid/kg) and the results are summarized in Figure 4.1. The level (u.g lipid/g tumour) of liposomal lipid in the LS180 and A431 tumours is shown in panels A and B , respectively, and the tissue concentration (ug drug/g tumour) of mitoxantrone in the L S I 8 0 and A431 tumours is shown in panels C and D , respectively. There are two important conclusions that can be made from the data shown in Figure 4.1, panels A and B . First, the accumulation rates of D M P C / C h o l and D S P C / C h o l liposomes are comparable in the L S I 8 0 tumours and they are comparable in the A431 tumours. Second, the rate of liposomal lipid accumulation in the L S I 8 0 tumour is significantly faster than that observed in the A431 tumour. In the L S I 8 0 tumour (Panel A ) the maximum concentration ( C m a x ) of liposomal l ipid observed is approximately 100-ug lipid/g tumour and at 4 hours following i.v. administration. In contrast, the C m a x of liposomal lipid observed in the A431 tumour (Panel B) is approximately 70-ug lipid/g tumour and at 48 hours after drug administration. Two important conclusions can be 88 inferred from the data shown in panels C and D. First, mitoxantrone accumulation in the solid tumours is increased when the drug is given encapsulated in liposomes in comparison to free drug. Fol lowing administration of free drug, the C m a x observed is at the 4 hour time point, a level of drug that is equivalent to that obtained following administration of the liposomal formulations of mitoxantrone. Subsequently the level of mitoxantrone observed in tumours decreases in animals given free mitoxantrone while the drug level increases or is maintained in tumours from animals given the liposomal formulations. Second, following administration of the liposomal formulations of mitoxantrone, the total concentration of drug achieved in the tumour is greater when drug is entrapped in D S P C / C h o l liposomes compared to D M P C / C h o l liposomes. This result is consistent with results from Chapter 3 demonstrating that the D M P C / C h o l liposomal formulation releases mitoxantrone more rapidly than DSPC/Cho l liposomes. Differences in the drug release attributes of these two liposomal formulations are emphasized in Figure 4.2 where the percentage of initial drug-to-lipid ratio is determined at the 48 hour time point. Panel A shows the percentage of initial drug-to-lipid ratio in the plasma compartment while panel B shows the percentage of initial drug-to-lipid ratio measured in isolated tumours. The plasma results are consistent with the results in Chapter 3, indicating that D S P C / C h o l liposomes retain 97% and 85% of the initial drug to lipid ratio in the plasma of mice bearing A431 or L S I 8 0 tumours, respectively. In contrast, the D M P C / C h o l formulations exhibit 22%> and 16% of the initial drug to lipid ratio in the plasma from mice bearing A431 or L S I 8 0 tumours, respectively. These values are comparable to those obtained in non-tumour bearing animals (see Chapter 3), suggesting that the presence of established tumours does not affect the release properties of the liposomes. Changes in drug to lipid ratios are less evident when the data are obtained from isolated tumours; however, these results (Panel B) are consistent with the plasma data and demonstrate a greater reduction in drug-to-lipid ratio for the D M P C / C h o l 89 liposomal mitoxantrone formulation. Data from A431 and L S I 8 0 tumours obtained from animals injected with the D S P C / C h o l formulation suggest that 90% and 78% of the entrapped mitoxantrone is still associated with the liposome, respectively. Tumours from animals injected with the D M P C / C h o l formulation have 53% (A431) and 62% (LS180) o f the drug associated with the liposome. It should be noted that the estimates of drug-to-lipid ratio in tumours are equivocal considering free mitoxantrone (or mitoxantrone that has been released from liposomes) w i l l localize in these regions of tumour growth (Figure 4.1, panel C and D). The extent of drug exposure in the two tumours is best summarized by the data in Table 4.1, which provides the area under the liposomal-lipid ( A U C L ) and mitoxantrone ( A U C D ) concentration-time curve values obtained in tumours from 0 to 96 hours following i.v. administration of free and liposomal drug (10 mg/kg drug dose). In the L S I 8 0 tumours, A U C L values of 10167 and 9926 ug of lipid/g of tumour/hour were measured following administration of mitoxantrone encapsulated in D S P C / C h o l and D M P C / C h o l liposomes, respectively. In the A431 tumour model, the D S P C / C h o l and D M P C / C h o l have A U C L values of 5728 and 5150 ug of lipid/g of tumour/hour, respectively. A comparison of the tumour A U C D values obtained after administration of the two liposomal formulations demonstrates that more drug is delivered using the D S P C / C h o l formulation (504 and 1000 ug drug/g of tumour/hour for the A431 and L S I 8 0 tumours, respectively) as compared to the D M P C / C h o l formulation (304 and 749 ug drug/g of tumour/hour for the A431 and L S I 8 0 tumours, respectively). It should be noted that the tumour A U C D values obtained after administration of free mitoxantrone are only 3 to 5 times lower than that measured for the liposomal formulations. This is in contrast to the area under the mitoxantrone concentration-time curves obtained in plasma from 0 to 96 hours, where the plasma A U C D is 20- to 30-fold lower following i.v. administration of free in comparison to that measured following injection of the liposomal formulations. 90 Figure 4.1 Lipid and mitoxantrone accumulation in A431 and LS180 tumours in SCID/RAG-2 mice over a 96 hour time period S C I D / R A G - 2 mice were injected bilaterally with 2 x 106 A431 cells and 1 x 10 6 LS180 cells subcutaneously. Once the tumours reached a size of approximately 0.05-0.2 cm 3 , mice were injected with a 10 mg/kg drug dose of free mitoxantrone (A), D S P C / C h o l mitoxantrone ( • ) , or D M P C / C h o l mitoxantrone ( • ) via the lateral tail vein. Mice were terminated using C 0 2 asphyxiation and tumours were removed and processed as described in Chapter 2, section 10. Panels A and B demonstrate lipid accumulation in both the L S I 8 0 and A431 tumours respectively and Panels C and D demonstrate drug accumulation in the L S I 8 0 and A431 tumours respectively. Data points represent the average and standard error o f the mean of at least 4 animals. Z2 O E -*—' O D) O E o 13) CD 160 120 24 48 72 Time (Hours) Time (Hours) 96 91 Figure 4.2 Percentage of initial drug to lipid ratio of DSPC/Chol mitoxantrone and DMPC/Chol mitoxantrone after 48 hours in plasma. S C I D / R A G - 2 mice were injected bilaterally with 2 x 10 6 A431 cells or 1 x 10 6 LS180 cells subcutaneously. Once the tumours reached a size of approximately 0.05 c m 3 , mice were injected with a 10 mg/kg drug dose of D S P C / C h o l mitoxantrone (shaded bars) or D M P C / C h o l mitoxantrone (open bars) via the lateral tail vein. Plasma and tumours were collected and processed as outlined in Chapter 2, section 10. Panel A shows the drug-to-lipid ratio in plasma and Panel B shows the drug-to-lipid ratio in the tumour. Data points represent the average and standard error of the mean of the data collected from at least 4 animals. A431 LS180 Tumours Tumours The distribution of drug from the plasma compartment to the tumour site can be described employing a drug targeting efficiency parameter, T e , relating the A U C in the circulation to the tumour A U C (T e = A U C T / A U C P ) . Using this parameter (see Table 4.1) it can be suggested that drug accumulation is more efficient in the L S I 80 tumours, an observation that is consistent with this tumour's extensive vascularization. The T e value obtained for the L S I 8 0 tumour is 2.3- to 2.8-fold greater than that observed for the A431 tumour. The T e values for the D S P C / C h o l and D M P C / C h o l liposomal mitoxantrone formulations are comparable for each tumour type and the greatest T e values obtained are for the free drug, and these values are at least 8-fold higher than those obtained for either liposomal formulation. This higher T e value for free drug is a reflection of drug distribution characteristics associated with small molecules (free drug) in comparison to large molecules (liposomal drug). 4.2.2 Efficacy of single dose administration of liposomal and free mitoxantrone in established A431 and L S I 8 0 human solid tumours In Chapter 3, the studies demonstrated that treatment of mice bearing L1210 and P388 liver tumours with D M P C / C h o l liposomal mitoxantrone resulted in 100% long term survivors. Although the D S P C / C h o l liposomal mitoxantrone formulation delivered more mitoxantrone than the D M P C / C h o l formulation to the tumour site, treatment with this formulation proved to be less effective due to drug release characteristics. It is important to determine whether these carrier-associated differences in mitoxantrone efficacy extend to solid tumours. A s indicated in the previous section, the A431 and L S I 8 0 tumours provided suitably different liposome uptake characteristics so that comparisons between the liposomal formulations could be made. It is important, however, to recognize that the selected tumour cells exhibit different growth characteristics and drug sensitivity (Table 4.2). Particularly, the L S I 8 0 tumours exhibit a growth 93 Table 4.1 A r e a under the liposomal-lipid and mitoxantrone concentration-time curves obtained in tumours and plasma from 0 to 96 hours following Iv. administration of free and liposomal drug (10 mg/kg dose) in S C I D / R A G - 2 mice bearing established A431 and LS180 tumours. Tumour L i p i d ug lipid/g of tumour/hour D S P C / C h o l D M P C / C h o l Drug ug drug/a of tumour/hour D S P C / C h o l D M P C / C h o l Free A431 L S I 80 5728 10167 5149 9925 505 1000 304 93 749 251 Drue ue drug/ml of plasma/hour Targeting efficiency (Te)a Plasma D S P C / C h o l D M P C / C h o l Free D S P C / C h o l D M P C / C h o l Free A431 373 226 11.6 1.35 1.35 8.01 LS180 254 264 11.8 3.94 2.84 21.3 " Targeting efficiency is a term that has been developed to characterize the distribution of drug between the plasma compartment and the tumour site. It is calculated by relating the A U C in the plasma compartment to the tumour A U C (T e = A U C T / A U C P ) . rate that is approximately 2-times faster than that measured for the A431 tumours. In contrast to the A431 tumours, L S I 8 0 tumours are highly vascularized and the L S I 8 0 cells are about 5-times more sensitive to free mitoxantrone in comparison to A431 cells. Gross observations indicated that the L S I 8 0 tumours are less cohesive than the A431 tumour and the L S I 8 0 tumours ulcerated more rapidly than A431 tumours. Results obtained following treatment of mice with established L S I 8 0 and A431 tumours are summarized in Figure 4.3. A s shown in Panel A , free mitoxantrone and the D S P C / C h o l 94 mitoxantrone formulation demonstrate minimal effects on the L S I 8 0 tumours. Tumour growth in animals treated with these formulations could not be distinguished from untreated controls, other than perhaps a reduction in the rate of tumour ulceration observed when animals were treated with free mitoxantrone. Animals that developed ulcerated tumours were killed as required by the Canadian Council for Animal Care guidelines and D S P C / C h o l mitoxantrone and saline treated animals typically exhibited tumour ulcerations when the volume exceeded 0.5 cm 3 . Reductions in tumour growth were observed when L S I 8 0 tumour bearing animals were treated with the D M P C / C h o l mitoxantrone formulation. It should be noted that treatment with this formulation did not result in a reduction in tumour size and the tumour growth rate measured after day 17 was equivalent to that observed for control mice. Although the L S I 80 cells are more sensitive to mitoxantrone than A431 cells in vitro (see Table 4.2) and L S I 8 0 tumours exhibited increased drug exposure (see Table 4.1) in comparison to the A431 tumours, the A431 tumours were more responsive to treatment with free mitoxantrone (Figure 4.3, Panel B) . Control mice exhibited 0.5 c m 3 tumours 12 days after initiation of treatment, whereas mice treated with free mitoxantrone exhibited similar tumour sizes after 16 days. The therapeutic activity of the liposomal formulations was better than free drug; however, there were slight differences in the therapeutic activity measured between liposomal formulations in the A431 tumours. M i c e treated with D S P C / C h o l liposomal mitoxantrone exhibited 0.5 cm 3 tumours 18 days after initiation of treatment versus 21 days with the D M P C / C h o l liposomal mitoxantrone. 4.2.3 Efficacy of multiple dose administration of liposomal and free mitoxantrone in non-established A431 and L S I 8 0 human solid tumours The studies summarized in Figure 4.3 were obtained when mice with well established tumours were treated with the different mitoxantrone formulations. It can be argued that optimal therapy 95 Figure 4.3 Efficacy of DSPC/Chol mitoxantrone, DMPC/Chol mitoxantrone and free mitoxantrone in established LS180 and A431 solid tumours in SCID/RAG-2 Mice S C I D / R A G - 2 mice were injected bilaterally with 1 x 10 6 L S I 8 0 cells (Panel A ) or 2 x 10 6 A431 cells (Panel B ) subcutaneously. Fourteen days after tumour cell inoculation (tumour size of > 0.05 cm 3 ) , mice were injected with a 5 mg/kg dose of free mitoxantrone (A), 10 mg/kg drug dose of D S P C / C h o l mitoxantrone ( • ) , or 10 mg/kg drug dose of D M P C / C h o l mitoxantrone ( • ) via the lateral tail vein. Control mice were injected with saline (V) . Tumour width and length were measured using calipers and volume was calculated as outlined in Chapter 2, section 14. Points represent average data and the standard error of the mean from at least 6 tumours. 96 Table 4.2 Attributes of the LS180 and A431 cell lines and their growth characteristics in SCID/RAG-2 mice. L S I 8 0 cells A431 cells Source Colon Carcinoma Squamous Ce l l Carcinoma Drug Sensitivity (in vitro)' Doxorubicin Vincristine Mitoxantrone .a 99 n M 14 n M 50 n M 83 n M 3 n M 275 n M Growth Rate (in vivo)b Characteristics 0.13 cnrVday Highly vascularized, poorly metastatic, loosely cohesive, mucin expressing 0.07 cmVday Poorly vascularized, metastatic, cohesive, E G F receptor positive and V E G F producing " Data refers to I C 5 0 concentrations, concentrations of drug that effects 50% growth inhibition or toxicity, determined in vitro during a 3-day continuous exposure cytotoxicity assay. Cell viability was determined using the M T T assay as described in the Chapter 2, section 7. b Growth rate was determined for control (untreated) tumours after the size exceeded 0.3 cm 3, a time point where significant increases in tumour size where measurable on a daily basis. should be observed when treating tumours at a time point prior to formation of a measurable tumour and through use of repeated injections of the drug. To address this, mice were treated with single and multiple doses of free and D M P C / C h o l liposomal mitoxantrone two days after tumour cell inoculation. The results of these studies have been summarized in Table 4.3. For simplicity the table reports results as the day of initiation of tumour growth, a parameter determined by taking a linear least-squares analysis of tumour volumes during the rapid growth phase and extrapolating to a tumour volume of zero. The effect of mitoxantrone treatment can then be determined as a delay in initiation of tumour growth. This analysis relies on the assumption that 97 Table 4.3 Treatment of non-established A431 and LS180 subcutaneous human xenografts in SCID/RAG-2 mice. Treatment is measured by estimations in the Delay in Tumour Growth" Initiation. Dose Schedule Treatment Dose Day of Delay in (mg/kg) Tumour Growth Tumour Growth (Days) LSI 80 tumours Control - 11 -Day 2 Free Mi tox 5 15 4 D M P C / C h o l Mi tox 10 15 4 Control - 12 -Days 2, 3, and 4 Free Mi tox 1.5 15 3 D M P C / C h o l Mi tox 3.5* 23 11 A431 tumours Control - 11 -Day 2 Free Mi tox 5 11 0 D M P C / C h o l Mi tox 10 15 4 Control - 17 -Days 2, 3, and 4 Free Mi tox 1.5 21 4 D M P C / C h o l Mi tox 3.5 27 10 " Determined as the day of initiation of tumour growth, a parameter determined by taking a linear least-squares analysis of tumour volumes during the rapid growth phase and extrapolating to a tumour volume of zero. It should be noted that treatment with mitoxantrone (free or liposomal) did not change the tumour growth rates, rather treatment effected a delay in the time when tumour growth initiated. * One mouse died due to toxic effects. treatment does not alter the growth rate of the tumour once it is established (i.e. tumour volume in excess of 0.05 cm 3 is attained). Treatment of non-established tumours with a single injection of D M P C / C h o l liposomal mitoxantrone at the maximum tolerated dose did not produce significant delays in tumour growth for the A431 and L S I 8 0 tumours. Following a single dose of free mitoxantrone, better therapeutic response was observed for mice bearing L S 180 tumours, where delays in tumour 98 growth o f 4 days were observed versus no delay in the A431 tumours. Although a number of different doses schedules were studied, including injections on day 2, 4 and 6 as well as day 2, 6 and 10, optimal therapy was observed for the day 2, 3, and 4 injection schedule reported in Table 4.3. Using this dose schedule, delays in A431 tumour growth of 4 and 11 days were obtained when mice were treated with free and D M P C / C h o l liposomal mitoxantrone, respectively. Delays in L S 180 tumour growth o f 3 and 11 days were obtained when mice were treated (day 2, 3 and 4) with free and D M P C / C h o l liposomal mitoxantrone, respectively. In all studies completed, the D M P C / C h o l liposomal mitoxantrone formulation was more active than free drug. A comparison of the D M P C / C h o l and D S P C / C h o l mitoxantrone formulation was made using the more sensitive L S I 8 0 tumour model and these results have been summarized in Table 4.4. These data support the conclusion that, regardless of dosing schedule or L S I 8 0 tumour burden, the D M P C / C h o l formulation of mitoxantrone is therapeutically more active than the D S P C / C h o l formulation. 4.2.4 Drug accumulation in the region of tumour cell inoculation The studies leading to the results summarized in Tables 4.3 and 4.4 raise an important question: Is a liposome-mediated increase in drug delivery achieved at the site o f tumour cell inoculation (i.e. prior to significant tumour growth)? This is a relevant question considering that the primary rationale used in the development of liposomal drug formulations is based on observations that demonstrate liposome-mediated increases in drug delivery to established tumours (see Figure 1). This observation has been attributed to the presence of blood vessels that are hyper-permeable to macromolecules in the plasma compartment and it is unlikely that such a vascular structure exists at a time point prior to significant tumour growth. In order to address this question, mitoxantrone delivery to the site of tumour cell inoculation was measured as described in Chapter 2, section 12. 99 Table 4.4 Treatment of non-established LS180 subcutaneous human xenografts in SCID/RAG-2 mice. Treatment is measured by estimations in the Delay in Tumour Growth" Initiation. Treatment* Dose (mg/kg) Day of Tumour Growth Delay in Tumour Growth (Days) LS180 tumours Control - 12 -Free Mi tox 1.5 14 2 D M P C / C h o l Mi tox 2.5 19 7 D S P C / C h o l Mi tox 2.5 14 2 " Determined as the day of initiation of tumour growth, a parameter determined by taking a linear least-squares analysis of tumour volumes during the rapid growth phase and extrapolating to a tumour volume of zero. It should be noted that treatment with mitoxantrone (free or liposomal) did not change the tumour growth rates, rather treatment effected a delay in the time when tumour growth initiated. * Mitoxantrone was administered i.v. on days 2, 6 and 10 after tumour cell inoculation Using a single time point (24 hours after drug administration), drug levels were measured in an area that included and surrounded the site of tumour cell inoculation. To confirm the presence of tumour cells in the site, mice were inoculated with radiolabeled L S I 8 0 cells and two days later the injection site was removed. Up to 75% of the injected radioactivity at the injection site could be recovered using this approach. It is recognized that this radioactivity can not be used as an indicator of cell number. Figure 4.4 shows the amount of mitoxantrone recovered at the site of tumour cell inoculation in comparison to drug levels measured (at the same time point) in established tumours. Although one set of results is obtained from tissue consisting primarily of tumour cells and associated host cells while the other consists of skin and muscle tissue, it does highlight two important points. 100 Figure 4.4 Drug accumulation at the site of tumour cell inoculation following i.v. administration of free mitoxantrone or mitoxantrone encapsulated in DMPC/Chol or DSPC/Chol liposomes S C I D / R A G - 2 mice were injected bilaterally with 1 x 106 LSI80 cells. Mice with established tumours [tumours with a volume > 0.05 cm 3 , (shaded bars)] or non-established tumours [mice treated 48 hours after tumour cell inoculation, (open bars)] were injected i.v. with a 10 mg/kg drug dose of D S P C / C h o l mitoxantrone, D M P C / C h o l mitoxantrone, and free mitoxantrone. [ , 4 C] labeled mitoxantrone was used as a tracer. 24 hours after treatment, established tumours were harvested and for non-established tumours, a 1.5 cm x 1.5 cm area surrounding the tumour cell injection site was harvested. Tissue was processed as described in Chapter 2. Data shown is the average of at least 6 tumours + the standard error. For comparison drug accumulation in established tumours is provided and these data were obtained from the data set used to generate Figure 1. 10 6H cn 3 o 10 DSPC DMPC Free 101 First, following i.v. administration of the liposomal mitoxantrone formulations the level of drug obtained in established L S I 8 0 tumours is 3-to 4-fold greater than that observed at the site of tumour cell inoculation. This difference in delivery is not observed following administration of free drug. Second, there is a 6-to7-fold increase in mitoxantrone delivered to established tumours when administering either liposomal formulation compared to free drug, however this difference decreases to less than 2-fold i f the injection site is evaluated 2 days following tumour cell injection. Since it is established that liposome accumulation in muscle tissue is typically undetectable, it can be suggested that drug delivery to the site of tumour cell inoculation is a consequence of liposome accumulation in the skin (Hwang et al., 1987; Gabizon et al., 1990; Yuan et al, 1994). 4.3 Discussion A central hypothesis that is guiding the development of lipid-based anti-cancer drug delivery systems in this thesis is that drug release is the most important attribute controlling the therapeutic benefits linked to use of liposomal carriers. Drug release is, of course, an ill-defined term that must take into account the rate at which a drug leaves the liposome. Depending on the drug encapsulated, slow drug release may foster decreases in drug toxicity (Mayer et al, 1989) and/or increases in therapeutic activity (Boman et al, 1994). Slow drug release has, however, also been linked to reduced drug biological availability and an associated decrease in anti-tumour activity (Mayer et al, 1989). Using mitoxantrone as an example, these studies have demonstrated that a slow drug release rate can effect a significant reduction in anti-tumour activity compared to faster-releasing carriers designed to exhibit comparable liposomal plasma elimination rates (L im et al, 1997). This conclusion was reached by comparing the anti-tumour 102 activity of mitoxantrone encapsulated in D S P C / C h o l and D M P C / C h o l liposomes following i.v. administration to mice bearing tumours residing primarily in the liver and spleen. The studies summarized in this chapter were initiated because of concerns that this conclusion was only applicable to liver localized disease, a site where significant and rapid accumulation of liposomes is observed following i.v. administration. In order to address this concern, the therapeutic activity of D S P C / C h o l and D M P C / C h o l liposomal mitoxantrone was measured using two human xenograft models grown as ectopic (s.c.) tumours. The results are considered by focusing this discussion on three important points, all critical i f the central hypothesis is to be sustained, including (1) the role of liposome delivery and liposome tumour/host cell interactions, (2) differences in drug targeting efficiencies between free and liposomal drug, and (3) the importance of considering capillary endothelium permeability to circulating macromolecules as well as capillary density within a tumour. For the anti-cancer drug doxorubicin, where benefits attributed to liposome delivery have been correlated to reductions in cardiotoxicity, reductions in the rate of doxorubicin release have been directly associated with reduced drug accumulation in cardiac tissue. Hence doxorubicin encapsulated in D S P C / C h o l liposomes is less toxic than doxorubicin encapsulated in egg P C / C h o i liposomes or D M P C / C h o l liposomes (Mayer et al, 1994). Interestingly, D M P C / C h o l liposomal doxorubicin, which releases 90% of its entrapped drug at a constant rate during the first 24 hours following i.v. injection, is approximately 3 times more toxic than free doxorubicin and more than 16 times more toxic than D S P C / C h o l liposomal doxorubicin, which releases less than 10% of its entrapped drug in vivo in the same period of time (Mayer et al, 1994). The rates o f drug release o f mitoxantrone are comparable to doxorubicin for these two liposomal lipid compositions. It is noteworthy that the maximum tolerated dose ( M T D ) of the D S P C / C h o l and 103 D M P C / C h o l mitoxantrone are comparable. In B D F 1 mice, the M T D of these formulations when given as a single i.v. injection was 20 mg/kg mitoxantrone (200 mg/kg lipid). In this study, which used S C I D / R A G - 2 mice, the M T D of these formulations (single i.v. dose) was 10 mg/kg mitoxantrone (100 mg/kg lipid). In contrast to doxorubicin formulated in D S P C / C h o l and D M P C / C h o l liposomes, both liposomal formulations of mitoxantrone were about half as toxic as free drug. This is an important point considering that the data shown in Figure 4.1 were collected following administration of free and liposomal mitoxantrone at 10 mg/kg. The free drug data were, therefore, obtained at a drug dose that would be toxic within a. 30-day time period and the resulting A U C D values are presumably an overestimate relative to the M T D of 5 mg/kg used for therapeutic studies. The 3-fold increase in drug exposure achieved using liposomal formulations o f mitoxantrone (Table 4.2) resulted in improvements in anti-tumour effects (see Figure 4.3). However, the results presented in this chapter do not support the notion that the greatest therapeutic activity w i l l be obtained using liposome formulations that facilitate the greatest increase in tumour drug A U C . The A U C D values obtained following administration of D M P C / C h o l mitoxantrone were 0.6 and 0.75 of the values obtained for D S P C / C h o l for the A431 and L S I 80 tumours, respectively. The D M P C / C h o l liposomal mitoxantrone formulation was therapeutically better than the D S P C / C h o l formulation when treating L S I 80 tumours (Figure 4.3A). Treatment of A431 tumours suggested that the D M P C / C h o l was as active as the D S P C / C h o l formulation (Figure 4.3B). Drug A U C values in solid tumours are dependent on the dose of lipid, the liposome plasma elimination rate as well as the drug retention characteristics of the liposome. The latter is illustrated by the data shown in Figure 4.1, where it is demonstrated that comparable liposomal lipid accumulation does not result in comparable drug uptake levels. In this example, reduction in 104 mitoxantrone uptake is partially a consequence of drug release from the D M P C / C h o l liposomes. This, however, is a simplistic analysis that does not account for the accumulation of drug released from liposomes in the plasma compartment or from other tissues that are accumulating and metabolizing liposomes. It has been proposed, for example, that the liver is capable of acting as a drug reservoir where macrophage processing of drug loaded liposomes can result in drug release back into the circulation (Storm et al, 1988). Indications of free (released) drug accumulation in tumours following i.v. administration of a liposomal drug have been based on comparisons between the estimated drug-to-lipid ratio in the plasma compartment versus the tumour. A s shown in Figure 4.2, the ratio of tumourd r u g.,0.]jpjd ratio/plasmadrug-to-iipid rat io at the 48 hour time point following administration of the D S P C / C h o l mitoxantrone is approximately 0.92 for both tumours. A similar analysis for the D M P C / C h o l mitoxantrone formulation results in a ratio of 2.5 for A431 tumours and 3.8 for L S I 80 tumours. A ratio of greater than 1 suggests that more drug is present in the tissue than would be predicted on the basis o f liposome accumulation from the plasma. The higher ratios observed in tumours following administration of D M P C / C h o l mitoxantrone are most likely a consequence of released drug accumulation. This can be suggested on the basis of the targeting efficiency (T e ) parameter, a value that is determined by dividing the A U C D in the tumour by the A U C D in the plasma compartment (see Table 4.1). The T e value for free mitoxantrone is at least 8-fold greater than that measured for the liposomal formulations. This is a consequence of differences in size between free drug and the liposomal drug. The free drug is small and readily distributes following i.v. administration, hence, the T e for free drug is large. Since drug is released from D M P C / C h o l liposomes while in the plasma compartment it is reasonable to assume that this drug could be efficiently taken into the tumour. It was demonstrated in this chapter that the rate and extent of liposome accumulation in tumours w i l l also be dependent on the type of tumour and this wi l l likely be a function of the tumour-105 specific attributes such as capillary density and structure. In the L S I 8 0 tumour model, liposome extravasation occurs rapidly, reaching the C m a x within four hours after administration (Figure 4.2A). In contrast, in the A431 tumour model the C m a x is achieved 48 hours after i.v. administration. Almost twice the amount of the liposomal lipid accumulates within the L S I 80 tumour ( A U C values of 10167.32 and 9925.82 ug lipid/g of tumour/hour for the D S P C / C h o l and D M P C / C h o l formulations, respectively) in comparison to the A431 tumours ( A U C values of 5728.22 and 5149.66 ug lipid/g of tumour/hour for the D S P C / C h o l and D M P C / C h o l formulations, respectively). Gross inspection of the tumours suggests that the L S I 80 tumour is better vascularized than A431 tumours (Table 4.2) and this may account for differences in rate of accumulation. It can be suggested that liposome extravasation may be dependent on tumour microvascular density as well as capillary endothelium permeability. The increased microvascular density would lead to greater delivery of liposomes to the site of tumour growth. In addition, the extravasation of liposomes is dependent on the permeability of the blood vessel. A n increase in the permeability (due to secreted factors such as V E G F ) could also result in increased liposome accumulation. A discussion relating microvascular density and endothelium permeability invites consideration of whether liposome extravasation is a relevant parameter when studying tumours prior to establishment of a significant tumour burden. For the L S I 80 and A431 tumours studied in this chapter, measurable tumours were obtained 12 to 15 days after tumour cell inoculation. It would be unexpected to see significant vascularization of the tumours shortly after cell inoculation, although no direct measurement of tumour vascularization was made in these studies. It can be suggested from the data shown in Figure 4.4 that liposome extravasation was reduced when the tumour burden was small. This is an indirect measurement and it should be noted that the results in Figure 4.4 compare drug accumulation in a tumour that has been carefully dissected from the 106 animal to drug levels measured in a large area of tissue that includes the cell inoculation site (as confirmed by recovery of radiolabeled cells) as well as surrounding skin and underlying muscle. Clearly it is important to develop methodologies that can measure liposomal lipid and drug levels in areas where tumour growth is initiating, particularly when considering that most studies evaluating liposome extravasation use large tumours that may have the greatest microvascular density and the most permeable blood vessels. It is also worth noting, however, that extravasation of liposomes into the peritoneal cavity in the absence of disease has been reported and this extravasation process is thought to be across normal vascular endothelium (Bally et al, 1993). A fundamental element of the central hypothesis is that drug encapsulated inside the liposome is not biologically available. Further, the Iiposome-encapsulated drug is not therapeutically active unless a feature promoting tumour cell delivery is incorporated. This may involve use of targeting ligands that are known to be internalized, for example the folate acid receptor (Lee and L o w 1993; 1994; Wang et al, 1995). In addition, non-internalized targets have also been used in an effort to specifically deliver the drug to tumour and release drug in the vicinity of the tumour cells (Longman et al, 1995; Scherphof et al, 1997). Alternatively, the liposomes can be designed to non-specifically bind and fuse with cells following extravasation into a site of tumour growth. A n elegant example of this approach, resulting in a lipid-based delivery system referred to as programmable fusogenic liposomes or P F V s , has recently been described (Holland et al, 1996). In the absence of cell delivery, cell fusion, and/or intracellular processing by phagocytic cells in the site of extravasation; however, the encapsulated drug must be released from the liposomes in order to maximize drug biological availability and therapeutic activity. 107 In conclusion, in order to fully maximize the benefits of using liposomal carriers, a balance between delivery and drug release must be achieved. It has been argued that the primary source of drug within the tumour is from liposomes that have extravasated into the site (Mayer et al, 1994), an argument that links the rate and extent of liposome accumulation and the rate of drug release to therapeutic activity. However, the possibility that drug release from sites that are distinct from the tumour may contribute to the therapeutic activity can not be excluded. This is perhaps most important when the tumour burden is small and vascularization is low. The results suggest that a conventional (non-targeted, non-fusogenic) formulation o f mitoxantrone prepared using D M P C / C h o l liposomes is active in treatment of ectopic (s.c.) tumours as wel l as tumours progressing primarily in the liver and spleen (see Chapter 3). This activity is believed to be a consequence of the rate at which mitoxantrone is released from D M P C / C h o l liposomes. The D M P C / C h o l formulation of mitoxantrone is particularly well suited for treatment of tumours (or sites of tumour growth) where liposome accumulation is rapid. The next chapter w i l l focus on the anti-tumour effects of D M P C / C h o l mitoxantrone when used to treat cancer within the liver. 108 CHAPTER 5 ROLE OF KUPFFER CELLS AND LIPOSOME MEDIATED DRUG DELIVERY TO LIVER IN GOVERNING THE EFFICACY OF DMPC/CHOL LIPOSOMAL MITOXANTRONE USED TO TREAT LIVER LOCALIZED CANCER 5.1 Introduction One o f the primary reasons for developing a liposomal formulation of an anti-cancer drug is to increase drug exposure at a site of tumour growth. Evidence to support this reasoning has come from many studies documenting that the maximum drug concentration as well as the length of time tumour drug levels are maintained is increased when an anti-cancer drug is administered inside an appropriately designed liposomal carrier (Parr et al, 1997; Bal ly el al, 1994; Gabizon, 1992; Mayer et al, 1990). Using mice bearing murine or human s.c. tumours, as much as 10% of the injected liposomal drug can be measured in association with an established tumour (Parr et al, 1997). Similar results are shown in Chapter 4 when mitoxantrone levels were evaluated in human xenograft models following i.v. administration of a liposomal formulation of mitoxantrone. The administration of liposomal mitoxantrone resulted in tumour mitoxantrone areas under the curve ( A U C D ) that were 4-to 5-fold greater then that observed following injection of free mitoxantrone. This improved delivery has been attributed to the presence of tumour-associated blood vessels that are hyperpermeable to circulating macromolecules (Yuan et al, 1995; W u etal, 1993; Kohn etal, 1992). Tumour drug levels are, however, low in comparison to those that can be obtained in the liver following parenteral administration of a liposomal anti-cancer drug. It was established over 20 years ago that liposomes have a tendency to localize in sites containing fenestrated blood vessels and high levels o f associated tissue macrophages, such as the liver (Rahman et al, 1982; Hinkle et al, 1978; Caride, 1976). Investigators have shown that liver drug exposure, as measured by 109 A U C D , can also be at least 5-fold greater than that which can be achieved with free drug (Zou et al, 1993a). Higher drug levels and increased exposure of the liver would imply that liposomal anti-cancer drugs should be well suited for use in the treatment of liver cancer. This has, however, been difficult to demonstrate. Although there are exceptions (Asao et al, 1992; Gabizon et al, 1983), approaches to treat hepatocellular carcinoma that use liposome-based delivery systems have been clumsy. The methods range from a reliance on immune stimulation (Okuno et al, 1998; Asao et al, 1992), administration via the hepatic artery (Cay et al, 1997; Konno, et al, 1995), the use of liposomes designed to release contents after an external stimulus is provided (Zou et al, 1993b) or on the use of a model that is based on i.v. injection of M5076 cells, a cell line known to actively take up liposomes by phagocytosis (Yachi et al, 1996). There are many possible explanations for why liposomal anti-cancer drugs have not been more successful in treating liver cancer. This would include an inherent insensitivity or resistance to cytotoxic drugs in tumour cells that arise in or metastasize to the liver (Furuya et al, 1997). Alternatively the blood vessels that arise in liver localized disease in response to angiogenesis signals may be less abundant (Toyoda, et al. 1997) and may exhibit altered vascular permeability to circulating macromolecules that is dependent on the microenvironment where the cancer grows (Fukumura et al, 1997). The latter point emphasizes that in the case of anti-cancer drug delivery to the liver, regional and cellular distribution of the drug may be critical i f therapeutic activity is to be obtained. In Chapter 3, a therapeutically active liposomal formulation of mitoxantrone for the treatment o f liver localized disease is described. The murine model used in this study was based on i.v. administration of L1210 cells into immune competent BDF1 mice ( F l D B A 2 / C 5 7 - B L 6 crosses). The L1210 cells are non-phagocytic and are sensitive to cytotoxic drugs. They have been and 110 continue to be used for assessing the in vivo activity of anti-cancer drugs (Canti et al, 1998; Perchellet et al, 1997; Gabr et al, 1997; Noda et al, 1997). The results from studies reported here suggest that the therapeutic activity of liposomal mitoxantrone is unequaled by other liposomal anti-cancer formulations prepared using comparable methods. In particular, it is demonstrated that liposomal formulations of doxorubicin and vincristine are only marginally active in the L1210 i.v. tumour model, a model that can be effectively cured when treated with D M P C / C h o l liposomal mitoxantrone (See Chapter 3). Such results provide an opportunity to address questions about what factors are important when considering development of a liposomal anti-cancer drug for use in the treatment of liver cancer. More specifically, this chapter addressed how liposome delivery to the liver may effect therapy in the i.v. L1210 tumour model. Two strategies designed to decrease liposomal delivery to the liver were employed. The first uses polyethylene glycol (PEG)-modified lipids to decrease serum protein binding (Du et al, 1997; Yuda et al, 1996) and liposome-cell interactions (Du et al, 1997; Yuda et al, 1996). The second method employs the use of agents (such as clodronate or doxorubicin) known to eliminate or impair Kupffer cells (Daemen et al, 1995, Parr et al, 1993; Bal ly et al, 1990; Van Rooijen and Classen, 1989). The results suggest that the therapeutic activity of liposomal mitoxantrone used to treat liver localized cancer is not dependent on the presence of Kupffer cells. However, strategies that non-specifically inhibit liposome-cell interactions (e.g. use of liposomes with PEG-modif ied lipids) significantly inhibit the therapeutic benefits achieve with D M P C / C h o l liposomal mitoxantrone. 5.2 Results 5.2.1 Therapeutic activity of free and liposomal anti-cancer drugs given i.v. to mice bearing the L1210 i.v. tumour model The L1210 i.v. tumour model was used to evaluate the efficacy of mitoxantrone, vincristine and doxorubicin administered i.v. in free form or encapsulated in liposomes (Table 5.1) . In chapter 3, it was demonstrated that following i.v. injection of 104 L1210 cells, tumour development is most evident in the liver and the spleen. The results in Table 5.1 were obtained following a single injection at a drug dose that was either the maximum tolerated dose (free and D S P C / C h o l vincristine; free and D S P C / C h o l mitoxantrone, E P C / C h o l doxorubicin) or at the lowest drug dose required to give maximum therapeutic effect (free doxorubicin and D S P C / C h o l doxorubicin and D M P C / C h o l mitoxantrone). Untreated and empty liposome (EPC/Cho l or D S P C / C h o l liposomes with encapsulated citrate buffer and p H 7.5 H B S outside and administered at a lipid dose of 150 mg/kg total lipid) treated animals were terminated as a result of significant tumour related disease within 10 days. The mean of the median survival time (9.8 days) was determined by averaging the median survival time for studies completed in D B A 2 mice (vincristine and doxorubicin treated animals, median survival time of 9.5 days) and those completed in B D F 1 mice (mitoxantrone treated animals, median survival time of 10 days). The significant point that can be made from the data in Table 5.1 is that the therapeutic activity o f D M P C / C h o l liposomal mitoxantrone (100% survival on day 60) is unequaled by the other drugs even when given in liposomal form. This result must, however, be considered in light of four other observations. First, 24 hour cytotoxicity assays measuring the cytotoxic/cytostatic activity of the free drugs (Table 5.2) suggest that L1210 cells are most sensitive to free mitoxantrone. This is consistent with the in vivo results shown in Table 5.1, where free 112 Table 5.1 Therapeutic activity of free and liposomal formulations of doxorubicin, vincristine and mitoxantrone following a single iv. injection in mice bearing the L1210 Lv. tumours. Treatment Drug Dose (mg/kg) Median Survival Time (days) %ILSe % Survival Control (saline) 9.8 a N / A Control (EPC/Chol ) 11.5 b 17 0 Control (DSPC/Chol ) 10.5 b 7 0 Free Mitoxantrone 10 17.0C 73 0 D S P C / C h o l Mitoxantrone 20 25.0° 155 0 D M P C / C h o l Mitoxantrone 10 >60 c N D f 100 Free Doxorubicin 10 13.5 b 38 0 E P C / C h o l Doxorubicin 30 1 8 b.d 84 0 D S P C / C h o l Doxorubicin 30 1 3 b,d 33 0 Free Vincristine 2 1 0 b,d 2 0 D S P C / C h o l Vincristine 3 1 3 . 5 M 38 0 a Determined in DBA2 and BDF1 mice and the value is based on the mean of the median survival time (days) in these two strains. b Determined in DBA2 mice c Determined in BDF1 mice d Indicates median survival times from one experiment using an n of at least 5 animals e Percentage ILS (Increase in Life Span) Values were determined from mean survival times of treated and untreated control groups. If greater than 50% of the animals survived more than 60 days the ILS% was not determined f Can not be determined because more than half the animals survived past 60 days mitoxantrone effected a 76% increase in life span (%ILS) compared to 38% ILS and 2% ILS obtained following treatment with doxorubicin and vincristine, respectively. Second, liposomal vincristine and liposomal doxorubicin are very effective when used i.v, to treat animals with i.p. L1210 tumours (Mayer, et al., 1993; Mayer et ai, 1989). Treating animals carrying i.p. L1210 tumours with D S P C / C h o l liposomal vincristine, for example, can result in greater than 50% long 113 Table 5.2 IC5o of doxorubicin, vincristine, and mitoxantrone when incubated with L1210 cells for 24 hours.a Drug I C 5 0 ( nM) a Doxorubicin 820 Vincristine 70 Mitoxantrone 55 a I C 5 0 is defined, based on the M T T assay described in the Chapter 2, as the concentration of drug where cell growth and/or viability is 50% of that observed in control (drug) free cultures. term (>60 day) survival. Third, D S P C / C h o l liposomal mitoxantrone is less active than the D M P C / C h o l formulation, a result that has been attributed to differences in the drug release rates from these two liposomes (See Chapter 3). It is important to note that the E P C / C h o l doxorubicin formulation (Bally et al, 1990; Harasym et al, 1997) and the D S P C / C h o l liposomal vincristine preparation (Mayer et al, 1993) have also been characterized as formulations that support release of entrapped contents following i.v. administration. Fourth, the most significant difference between the liposomal formulations of vincristine, doxorubicin and mitoxantrone is that the vincristine and doxorubicin formulations induce hepatic M P S blockade (Daeman et al, 1995; Bal ly et al, 1990). It is for this reason that this chapter w i l l evaluate the influence of hepatic M P S avoidance and elimination strategies on the activity of the D M P C / C h o l mitoxantrone formulation. 5.2.2 Reducing D M P C / C h o l liposomal mitoxantrone delivery to the liver Two strategies were used to effect reductions in the delivery of D M P C / C h o l liposomal mitoxantrone to the liver. One involved incorporation of PEG-modified lipids into the 114 D M P C / C h o l formulation (hepatic M P S avoidance strategy) and the second involved administering a pre-dose of D S P C / C h o l doxorubicin (2 mg/kg drug) 24 hours prior to administration of D M P C / C h o l liposomal mitoxantrone (hepatic M P S elimination strategy). As illustrated in Figure 5.1A, 5.IB and 5.2A, it was anticipated that both strategies would cause a decrease in the rate of liposomal lipid (Fig. 5.1 A ) and mitoxantrone (Fig. 5. IB) elimination from the plasma compartment and an associated decrease in drug accumulation in the liver (Figure 5.2A). For example, 24 hours after i.v. administration of D M P C / C h o l mitoxantrone, the level of mitoxantrone measured (using a [ l 4C]-labeled drug as a marker) in the liver was 27 ug/g of liver. When mitoxantrone was administered in D M P C / C h o l liposomes with 5 mol % PEG2ooo-modified lipids the drug levels in the liver at 24 hours were reduced to 12.2 ug/g of liver. When the mice were given the pre-injection of D S P C / C h o l liposomal doxorubicin (2 mg/kg drug), mitoxantrone levels in the liver 24 hours after administration of D M P C / C h o l mitoxantrone were below 8 u.g/g of liver. The greater than two-fold reduction in liver mitoxantrone levels measured at 24 hours was associated with approximately a 3-fold and 5-fold increase in plasma concentrations of drug and liposomal lipid, respectively. The plasma elimination rates over the first 24 hours after administration were comparable for the PEG-containing liposomes and the D M P C / C h o l mitoxantrone formulations given to mice pre-injected with D S P C / C h o l liposomal doxorubicin. 5.2.3 Influence of reducing liver mitoxantrone levels on the therapeutic activity of D M P C / C h o l mitoxantrone Figure 5.2B demonstrates how the two strategies for reducing D M P C / C h o l mitoxantrone delivery to the liver affected its therapeutic activity when used to treat the L1210 i.v. tumour model. The results obtained were surprising. Incorporation of PEG2ooo->Tiodified lipids into the D M P C / C h o l mitoxantrone resulted in a significant reduction in therapeutic activity. In dramatic contrast, pre-115 Figure 5.1 Plasma Elimination of DMPC/Chol Mitoxantrone Liposomes and DMPC/Chol/PEG Mitoxantrone Liposomes. M i c e were pre-treated with 2 mg/kg drug dose of D S P C / C h o l Doxorubicin in order to induce M P S blockade. 24 hours later, M P S Blockade mice were injected with D M P C / C h o l mitoxantrone ( • ) . N o n - M P S Blockade mice were treated with 10 mg/kg dose of D M P C / C h o l mitoxantrone ( • ) or D M P C / C h o l / P E G mitoxantrone ( T ) . Blood was collected as described in Chapter 2. Panel A shows elimination of lipid from the plasma compartment over 24 hours. Panel B shows the elimination of drug from the plasma compartment over 24 hours. Points represent the average and the standard error of at least 8 mice. * signifies p<0.05. 0 4 8 12 16 20 24 Time (Hours) 116 Figure 5.2 Drug accumulation in the liver versus therapeutic activity In Panel A , drug delivery to the liver was assessed using l 4C-mitoxantrone as a tracer. CD1 mice were injected with a 10 mg/kg drug dose of D M P C / C h o l mitoxantrone. M P S blockade treated mice were injected with a 2 mg/kg drug dose of D S P C / C h o l Doxorubicin 24 hours prior. Livers were harvested and processed as described in Chapter 2. Bars represent the average and standard error collected from 8 mice. * symbolizes significant differences from the D M P C / C h o l mitoxantrone group (p < 0.05). In Panel B , therapeutic activity was assessed. B D F 1 mice were inoculated with 1 x 10 6 L1210 tumour cells. M P S blockade mice were treated two hours after tumour cell inoculation. 24 hours after tumour cell inoculation, mice were treated with a 10 mg/kg dose o f D M P C / C h o l mitoxantrone. Dashed line represents the survival time of untreated mice. ** indicates greater than 60 day survival. 35 30 25 20 15 10 60 50 .1 40 g 30 w ro 20 10 DMPC/Chol DMPC/Chol/PEG-PE DMPC/Chol Mitoxantrone Mitoxantrone Mitoxantrone + RES Blockade 117 treatment with D S P C / C h o l doxorubicin had no impact on the therapeutic activity o f the D M P C / C h o l mitoxantrone formulation. Further support of these data is provided in Table 5.3. The rationale for these studies is based on the potential that the pre-dose of D S P C / C h o l liposomal doxorubicin may have therapeutic activity. A s indicated in Table 5.1, this formulation has minimal activity (< 20 % ILS) when used to treat the L1210 i.v. tumour model at doses of 30 mg/kg. The activity of this formulation, however, could be augmented by mitoxantrone. In order to address this issue, two other approaches to achieve hepatic M P S blockade were used, including a pre-dose o f liposomal vincristine or liposomal clodronate. Although vincristine is also an anti-cancer agent, its mechanism of activity is distinct from doxorubicin. A s noted in Table 5.1, liposomal vincristine is also not active when treating the L1210 i.v. tumour model. Clodronate is a bisphosphonate that has been developed for treatment of osteoporosis (Fleisch, 1993; Lepore et al, 1991) and is known to deplete macrophages, particularly well when given in liposomal form (Van Rooijen, and Claassen, 1988; Van Rooijen and Van Nieuwmegen, 1984). In addition, the influence of hepatic M P S blockade, achieved using the three different pre-treatment strategies, on the therapeutic activity of the PEG-containing D M P C / C h o l mitoxantrone was assessed. The results presented in Table 5.3 are unambiguous. First, hepatic M P S blockade achieved by pre-treating animals with liposomal doxorubicin, vincristine or clodronate had no impact on the median survival time of mice bearing the i.v. L I210 tumours. Second, the therapeutic activity of the D M P C / C h o l / P E G mitoxantrone formulation was not affected by any of the pre-treatment strategies. Third, regardless of what agent was used to achieve hepatic M P S blockade, mice treated with D M P C / C h o l mitoxantrone exhibited 100% long term (>60 day) survival. 118 Table 5.3 Influence of PEG-lipid incorporation and hepatic MPS blockade on the L1210 Anti-tumour Activity of DMPC/Chol Mitoxantrone Pre-Treatmenta Treatmentb Median Survival %ILS C /o Time Survival None Untreated 9.5 0 Dox Blockade 9 - 0 Vine Blockade 11 16 0 Clodronate Blockade 9 - 0 None D M P C / C h o l Mi to > 60 days N . D . d 100 Dox Blockade (10 mg/kg) > 60 days N . D . 100 V ine Blockade > 60 days N . D . 100 Clodronate Blockade > 60 days N . D . 100 None D M P C / C h o l / P E G Mi to 17 79 Dox Blockade (10 mg/kg) 20 111 0 Vine Blockade 15 58 0 Clodronate Blockade 18.5 94 0 a Pre-treatment was administered two hours after tumour cell inoculation b Treatment dose was at drug dose of 10 mg/kg at drug to lipid ratio 0.1 (wt:wt) 0 Percentage Increase in Life Span (ILS) values were determined from median survival times of treated and untreated control groups. d Can not be determined because more than half the animals survived past 60 days 5.2.4 Influence of hepatic M P S avoidance and elimination strategies on mitoxantrone release In the previous chapters, it was postulated that the therapeutic activity of liposomal mitoxantrone is dependent on the rate of mitoxantrone release from the liposomes following administration. Therefore, it was important to determine whether the hepatic M P S avoidance and elimination strategies affected drug release rates. As shown in Figure 5.3, there was a significantly higher drug-to-lipid ratio observed at 24 hours following injection of D M P C / C h o l / P E G mitoxantrone in comparison to D M P C / C h o l mitoxantrone, suggesting that the drug release is inhibited in 119 Figure 5.3 Drug release of mitoxantrone from DMPC/Chol liposomes and DMPC/Chol/PEG liposomes M i c e were pre-treated with 2 mg/kg drug dose of D S P C / C h o l Doxorubicin in order to induce M P S blockade. 24 hours later, M P S Blockade mice were injected with D M P C / C h o l mitoxantrone ( • ) . N o n - M P S Blockade mice were treated with 10 mg/kg dose of D M P C / C h o l mitoxantrone ( • ) or D M P C / C h o l / P E G mitoxantrone ( T ) . Blood was collected as described in Chapter 2, Section 9. Points represent the average and the standard error of at least 8 mice. * signifies p<0.05. 03 CL C O =3 0.100 0.075 0.050 0.025 H 0.000 12 16 20 Time (Hours) 120 liposomes with the PEG-modified lipid. This was surprising considering results with vincristine suggest that drug release rates are increased when the liposomes used contain PEG-modif ied lipids (Webb et al, 1998). However, in this case the decrease in protein adsoprtion to the surface o f the liposome due to the addition of PEG-modified lipids may have a role in the increased retention of mitoxantrone. It is possible that reduced therapeutic activity is a consequence of reduced drug release from the D M P C / C h o l / P E G mitoxantrone formulation. It is important, however, to note that mitoxantrone release from D M P C / C h o l / P E G liposomes is faster than that observed for the D S P C / C h o l formulation and its therapeutic activity is less than that observed for D S P C / C h o l liposomes (Chang et al, 1997). The study reported by Chang et al. (1997) also provided data suggesting that the therapeutic activity of a P E G containing formulation was less than that observed for liposomes prepared in the absence of PEG-l ip ids . A s expected, strategies relying on the use of hepatic M P S blockade had no effect on drug release from the D M P C / C h o l liposomes (Fig. 5.3). 5.2.5 Influence of hepatic M P S avoidance and elimination strategies on liposome distribution in the liver and on Kupffer cell depletion Induction o f hepatic M P S blockade was achieved by injecting a low dose (2 mg/kg drug) of D S P C / C h o l doxorubicin and by the more established technique involving use of liposomal clodronate. Confirmation that these strategies caused depletion of Kupffer cells is provided in the micrographs shown in Figure 5.4. These micrographs were obtained by staining liver cryosections with an antibody (F4/80) that labels mature macrophages (Lee et al, 1985; Hume et al, 1984; Austyn and Gordon, 1981). Sections derived from livers of untreated mice (Panel A ) contain many F4/80 positive cells, cells that are presumed to be liver Kupffer cells. The population of labeled cells is reduced significantly when the liver sections are obtained from 121 mice that had been injected 24 hours earlier with liposomal doxorubicin (Panel B) or liposomal clodronate (Panel C) . The reduction in F4/80 positive cells was most significant in the clodronate treated animals. The data presented in Figure 5.4 is consistent with other reports (Van Rooijen et al, 1990) however it has not been established how macrophage depletion or macrophage avoidance ( P E G -liposomes) impacts the distribution of liposomal mitoxantrone in the liver. In order to obtain this information two approaches were taken. First, the liposomal mitoxantrone formulations, either D M P C / C h o l or D M P C / C h o l / P E G , were prepared with the fluorescent lipid l , l ' -d ioctadecyl-3,3,3',3'-tetramethylindocarbocyanine perchlorate (Dil) . It has been demonstrated that this fluorescent lipid does not exchange with neighboring membranes (Claassen, 1992; Honig and Hume, 1986) and thus it is considered as a useful marker for liposomes in vivo. Twenty-four hours following i.v. administration of D i l labeled D M P C / C h o l mitoxantrone and D M P C / C h o l / P E G mitoxantrone (10 mg/kg drug dose), livers were removed, crysections were prepared and the sections were viewed using confocal microscopy. As seen in Figure 5.5, incorporation of P E G modified lipids caused a reduction in liposome accumulation in the liver (compare panel A to panel B) . Changes in the distribution of D i l labeled D M P C / C h o l mitoxantrone in the liver are more dramatic in livers isolated from mice pre-treated with liposomal formulations of doxorubicin, vincristine or clodronate (Figure 6). Hepatic M P S blockade caused a significant reduction in the amount of fluorescently labeled D M P C / C h o l mitoxantrone delivered to the liver (compare Panel A to Panels B - D ) . In addition to the decrease in liposome accumulation, the liposome distribution pattern is changed considerably and the distribution pattern is different when comparing liposomal doxorubicin (Panel B) and vincristine (Panel C) induced hepatic M P S blockade to that observed with liposomal clodronate (Panel D) . 122 Figure 5.4 F4/80 staining of Kupffer cells in the liver Livers from CD1 mice were pre-treated with either D S P C / C h o l Doxorubicin or E P C / C h o l Clodronate. Control liver was left untreated. 24 hours later, livers were extracted and embedded in O.C.T . media. A s outlined in the Chapter 2, section 17, livers were then stained with the F4/80 antibody. Magnification is 40x for all panels. Arrows indicate stained Kupffer cells. 123 124 Figure 5.5 Confocal imaging of biodistribution of Dil labeled DMPC/Chol mitoxantrone liposomes and DMPC/Chol/PEG mitoxantrone liposomes in the liver M i c e were injected with a 10 mg/kg drug dose of D i l labeled D M P C / C h o l mitoxantrone or D M P C / C h o l / P E G mitoxantrone. 24 hours later, mice were terminated via C 0 2 asphyxiation, and livers harvested. Livers were processed as outlined in the Chapter 2, section 19 and imaged using a B ioRad 6000Z Confocal Imaging System. Panel A represents images from mice injected with D M P C / C h o I / D i l mitoxantrone and Panel B represents images from mice injected with D M P C / C h o l / P E G / D i l mitoxantrone. Magnification is lOx for all panels. 125 126 Figure 5.6 Confocal imaging of biodistribution of D i l labeled D M P C / C h o l mitoxantrone liposomes with and without M P S blockade. M P S blockaded mice were pre-treated with a 2 mg/kg drug dose of D S P C / C h o l doxorubicin or 1 mg/kg vincristine or EPC/Clodronate. N o n - M P S blockaded mice were left untreated. 24 hours later, mice were injected with a 10 mg/kg drug dose of D i l labeled D M P C / C h o l mitoxantrone. 24 hours after injection, mice were terminated via C 0 2 asphyxiation, and livers harvested. Livers were processed as outlined in Chapter 2, Section 19and imaged using a BioRad 6000Z Confocal Imaging System at lOx. Panel A represents images from non MPS-blockaded mice, Panel B are images from mice with M P S Blockade using D S P C / C h o l doxorubicin at a 2 mg/kg drug dose, Panel C are images from mice with M P S Blockade using D S P C / C h o l vincristine at a 1 mg/kg drug dose, and Panel D are from mice with M P S Blockade using E P C / C h o l clodronate. Magnification is lOx for all panels. 127 Fol lowing hepatic M P S blockade with liposomal doxorubicin and vincristine, D i l labeled D M P C / C h o l mitoxantrone distributed in discrete patches. Numerous vacuoles are seen in the micrographs of livers from liposomal vincristine pre-treated mice. These may attributed to vincristine induced autophagocytosis in hepatocytes and the associated appearance of autophagocytic vacuoles (Hirsimaki and Pilstrom, 1982). The distribution pattern observed in animals pretreated with liposomal clodronate (Panel D) is comparable to that observed for D i l labeled D M P C / C h o l mitoxantrone, except there are fewer liposomes present. Mitoxantrone delivery to liver hepatocytes was also measured in an effort to resolve differences between the D M P C / C h o l mitoxantrone (in the presence and absence of hepatic M P S blockade population) and the formulation prepared with PEG-modified lipids. Hepatocytes were isolated as described in Chapter 2 and the level of drug was measured using [ 1 4C]-mitoxantrone as a marker for drug. Drug levels were standardized to 106 hepatocytes. It should be noted that hepatocyte drug levels may be due, in part, to drug that has been taken up during the hepatocyte isolation procedure. Given this analysis, it was anticipated on the basis of the data presented in Figure 5.2A and Figure 5.4, where hepatic M P S blockade affected a 2- to 3-fold reduction in liver mitoxantrone levels and a significant (>90%) reduction in Kupffer cells, that hepatocyte delivery would increase significantly when M P S blockade was used. A s shown in Figure 5.7, this was not the case. Liposomal doxorubicin and clodronate pre-treatment effected a 2-fold reduction in liposome delivery to the hepatocytes, a reduction that is comparable to that observed in the whole liver. When hepatocyte mitoxantrone levels were determined in animals given (i.v.) D M P C / C h o l / P E G mitoxantrone the values also decreased by a factor of 2. It can be suggested that differences in the anti-tumour activity of D M P C / C h o l liposomal mitoxantrone due to P E G -lipid incorporation or hepatic M P S blockade can not be attributed to altered drug delivery to hepatocytes or to Kupffer cell processing. 128 Figure 5.7 D r u g delivery to hepatocytes Non-blockaded female CD1 mice were injected with a 10 mg/kg drug dose D M P C / C h o l mitoxantrone (A) or D M P C / C h o l / P E G mitoxantrone (D). M P S blockaded mice were pre-treated with either D S P C / C h o l Doxorubicin (B) or E P C / C h o l Clodronate (C). Twenty four hours later, the mice were then treated with 10 mg/kg drug dose D M P C / C h o l mitoxantrone. Livers were extracted and hepatocytes isolated as described in the Chapter 2, Section 18. L i p i d and drug concentrations were assessed via scintillation counting for 3 H and l 4 C . Bars represent the average + standard error. 0.12 £ 0.10 -i — ' — i o I 008 -X CD ° 0.06 -X . J- . "I 0.04 - _ r r ~ •o o) 0.02 -o.oo J — I — L - I — L J — I A B C D 129 5.3 Discussion There are two very simple conclusions that can be made on the basis of the data presented in this chapter. First, Kupffer cells do not play a role in governing the therapeutic activity of D M P C / C h o l liposomal mitoxantrone. Second, incorporation of PEG-modif ied lipids significantly inhibits the therapeutic activity of D M P C / C h o l liposomal mitoxantrone. The question that needs to be addressed on the basis of these conclusions is equally simple: Why should one strategy designed to reduce drug delivery to the liver inhibit therapy while another, which achieves a similar reduction in drug delivery, have no effect? To address this question it is important to examine the assumptions made when designing the experiments. These assumptions included: 1) drug delivery to the site of disease is critical in defining the therapeutic activity of liposomal mitoxantrone when used to treat liver localized disease; 2) conversely, reduction in drug delivery to the liver would effect reduced therapeutic activity and the related assumption 3) that PEG- l ip id mediated reductions in liver delivery would provide similar results when compared to strategies relying on use of hepatic M P S blockade. The three assumptions, in retrospect, seem quite naive. The first assumption that drug delivery is critical in defining the therapeutic activity has, in effect, been addressed by previous investigators and confirmed by results shown in Table 5.1. A s indicated in the introduction, many liposomal anti-cancer drugs have not been particularly effective in the treatment of liver cancer. This can be attributed to the role of the liver in drug metabolism and detoxification of drugs (Meijer et al, 1990; Erlinger, 1996; Yamazaki et al, 1996) and to inherent drug resistance of colon cancer and hepatocellular carcinomas (Ferry, 1998). It is believed that the latter concern is not really an issue in the present study because the cell line used (murine L1210 cells) was quite sensitive to the drugs selected (see Table 5.2). 130 Although the cytotoxicity assay would suggest that the L1210 cells are approximately 10-fold less sensitive to doxorubicin, it has been demonstrated that free doxorubicin and liposomal doxorubicin are quite effective in treating animals bearing L1210 tumours in the peritoneal cavity (Mayer et al. 1989). For this reason, it can be presumed that difference in therapeutic activity of these drugs, in free or liposomal form, are a consequence of differences in drug metabolism in the liver and elsewhere. If comparisons are restricted to the anthraquinone mitoxantrone and the anthracycline doxorubicin, then some critical determinants of activity can be identified. The most significant difference in these drugs concerns their ability to generate free-radicals. In the presence of rat liver microsomes and the electron donor N A D P H , doxorubicin is reduced to its free radical form and under identical conditions mitoxantrone is not (Vi le and Winterboum, 1989). The cytotoxic properties of doxorubicin have been attributed to generation of semi-quinone radicals that subsequently enter redox cycles with molecular oxygen which, in turn, lead to cation-radical formation (Riley and Hanzlik, 1994). This is associated with doxorubicin mediated stimulation of superoxide anion production that is not observed for mitoxantrone (Basra et al, 1985). It is believed that doxorubicin cardiotoxicity is mediated by free radical production and lipid peroxidation (Vi le and Winterbourn, 1989) and differences in generation of reactive oxygen have been used to explain why mitoxantrone exhibits reduced cardiotoxicity. The same argument has been used to explain why liposomal formulations of mitoxantrone do not promote hepatic M P S blockade (Chang et al, 1997; L i m et al, 1997) while formulations of doxorubicin are so effective in depleting non-dividing cells of the M P S (Bally et al, 1990; Daemen etal, 1995). Part of the rationale used in the hepatic M P S blockade studies was based on the fact that liposomal mitoxantrone does not induce M P S blockade, while liposomal vincristine and 131 liposomal doxorubicin do induce M P S blockade. Previous studies have suggested that Kupffer cells can play a role in processing liposomal anti-cancer drugs (Storm et al, 1988), providing a mechanism for drug release back into the systemic circulation and/or within the region o f macrophage localization. It was, therefore, convenient to suggest that the reason why liposomal formulations of doxorubicin and vincristine were not active in the treatment of liver localized disease related to hepatic M P S blockade. Conversely liposomal mitoxantrone activity is due, in part, to Kupffer cell processing. The data presented in Figure 5.2B and Table 5.3 clearly demonstrate that this is not the case. The therapeutic activity of liposomal mitoxantrone is not influenced under conditions where Kupffer cells have been eliminated. There are other attributes of mitoxantrone that may make it better suited for treatment o f liver localized cancer. For example, it is established that the cytotoxic activity of mitoxantrone is dependent on functional cytochrome P450-dependent mixed function oxidase (Duthie and Grant, 1989), a result that suggests that a mitoxantrone metabolite may be the primary effector of cytotoxicity (Mewes et al, 1993). Ramirez et al (1996) has argued that mitoxantrone may be a good agent for treatment of liver disease because its main route of metabolism is within the liver and using a hepatic tumour model in rabbits, this group demonstrated that hepatic artery administration of mitoxantrone provided better therapy then intravenous administration. These data were used to support the conclusion that regional administration o f mitoxantrone should be considered for treatment of liver cancer. Perhaps the properties of D M P C / C h o l liposomal mitoxantrone that facilitate increases in drug exposure account for the improved activity observed when the liposomal drug is given intravenously. It should be noted that there is a potential concern regarding the use of mitoxantrone to treat liver disease in mice. Schrenk et al. (1996) have suggested that mitoxantrone is not an efficient inducer of mdrl gene expression in 132 murine liver, which contrasts results obtained in rats. The mdrl gene encodes for an A B C transporter known to play a role in biliary excretion of certain xenobiotics (Schrenk et al, 1993). In the /.v. L1210 tumour model, the anti-tumour activity between different liposomal drugs can be accounted for by unique attributes of the drug used; however, it is difficult to explain differences between the D M P C / C h o l (in the presence and absence of hepatic M P S blockade) and the D M P C / C h o l / P E G formulations. Perhaps the most compelling argument is one based on P E G -mediated inhibition of liposome-cell interaction. Conversely, the therapeutic activity o f mitoxantrone is dependent on cell processing but the cells involved are not mature liver macrophages. The former argument is supported by data demonstrating that P E G modification inhibits protein and cell binding (Du et al, 1997). Inhibition of cell binding is observed even when targeting ligands are attached to the liposome surface (Harasym et al, 1995) and i f cell binding is obtained, the presence of PEG-modified lipids may prevent endocytosis (Ishiwata et al, 1997). In terms of the counter-argument, that cell processing is required for optimal therapeutic activity, it is essential to expand our discussion beyond the role of Kupffer cells. It is established that several cell types in the liver may be responsible for removal of particles from the blood compartment (Shiratori et al, 1993; Bouwens et al, 1992). Two populations of cells are of particular interest. Sinusoidal endothelial cells are capable of endocytosis and can accumulate particles <200 nm. In addition, Shiratori et al. (1993) have shown that when Kupffer cell function is blocked, sinusoidal endothelial cells can provide a compensating role in particle removal. The mechanism of particle removal by endothelial cells is believed to be different then that o f Kupffer cells (Dan and Wake, 1985). The second population of interest is monocytes (van Furth, 1980). Bouwens and Wisse (1985) have argued that there are two populations o f 133 phagoctes in the liver, a result that has been confirmed by more recent immunocytochemical analysis (Armbrust and Ramadori, 1996). Further, it has been demonstrated that there can be significant extrahepatic recruitment of monocyte-derived macrophage precursors in liver (Bouwens et al, 1986). The hepatic M P S blockade strategies employed in our studies may have been sufficient to eliminate Kupffer cells, but cell internalization and processing by monocytes that have been recruited to the liver, by immature liver phagocytic cells and by sinusoidal endothelial cells may all contribute to the activity of D M P C / C h o l mitoxantrone. Differences in the activity o f D M P C / C h o l mitoxantrone and the D S P C / C h o l formulation could still be attributed to drug release properties following cell uptake. It is important to note that one can not entirely eliminate the possibility that the reduced activity o f D M P C / C h o l / P E G mitoxantrone was due to reduced drug release rates (see Figure 5.3). The observation that the PEG-containing formulation released drug slower then the D M P C / C h o l formulation was surprising and was contrary to results obtained with liposomal vincristine (Webb et al, 1998). The latter observation was attributed to PEG-mediated changes at the membrane interface that could favor increased partitioning of the drug into the membrane. A n alternative model to explain the PEG-induced decreases in mitoxantrone release may involve the influence of serum protein binding on mitoxantrone release from the D M P C / C h o l liposomes. Consistent with the drug release argument, we can also not exclude the possibility that drug release from liposomes in the plasma compartment or from a site distinct from the liver may contribute to the therapeutic activity and that cell processing is not important. A s indicated in the results, however, it is believed that the rate of drug release from D M P C / C h o l / P E G mitoxantrone is sufficient to obtain therapy. This conclusion is based on results obtained with D S P C / C h o l formulations o f mitoxantrone that are more active in treating the i.v. L1210 tumour model despite having slower (See Chapter 3) or equivalent (Chang et al, 1997) drug release characteristics. 134 Therefore, it is concluded that reductions in therapy observed for D M P C / C h o l / P E G mitoxantrone were due to inhibition of cell binding and processing. Conversely the activity o f the D M P C / C h o l mitoxantrone is dependent on cell processing, but Kupffer cells do not play a significant role in this processing step. 135 C H A P T E R 6 S U M M A R I Z I N G D I S C U S S I O N 6.1 S u m m a r y of results The objective of the studies presented in this thesis was to outline the importance of drug release in the development of liposomal mitoxantrone. Drug release was evaluated in liver localized disease, a site where rapid liposome accumulation occurs. This was then extended to studies evaluating drug release at a site where liposome accumulation is slow, such as. a subcutaneous tumour. Finally, the activity of liposomal mitoxantrone was evaluated in the liver where the effects of drug delivery were assessed. In Chapter 3, the influence o f liposome drug release on the therapeutic activity o f encapsulated mitoxantrone was reported. In vivo studies demonstrated that D M P C / C h o l liposomes released drug faster than D S P C / C h o l liposomes. Efficacy studies were conducted in BDF1 mice inoculated i.v. with murine P388 cells or L1210 tumour cells. M i c e treated with a single dose of 10 mg drug/kg of D M P C / C h o l liposomal mitoxantrone resulted in 100% of the treated animals surviving for more than 60 days. In contrast, no long term survivors were obtained in any other treatment group, even when drug doses were escalated to the M T D . Pharmacodynamic studies with D M P C / C h o l mitoxantrone and D S P C / C h o l liposomal mitoxantrone illustrate the importance o f achieving a balance between drug release characteristics and drug delivery to a site o f tumour progression. In Chapter 4, delivery and therapeutic activity of liposomal mitoxantrone formulations exhibiting different drug release characteristics in two human carcinoma xenograft models (A431 and 136 LSI80) that accumulate liposomes at different rates was evulated. When lipid and drug levels were measured in established (> 0.05 cm 3 ) tumours, accumulation was more rapid in the L S I 8 0 tumours ( C m a x 4 hours) when compared to the A431 tumours ( C m a x 48 hours). A U C values for liposomal lipid measured over a 96 hour time course were comparable for both liposomal formulations in A431 and the L S I 8 0 tumours, however liposomal lipid A U C values were almost 2-fold higher in L S I 8 0 tumours than in A431 tumours. Although drug delivery was less following administration of the D M P C / C h o l liposomal mitoxantrone in comparison to the D S P C / C h o l formulation, anti-tumour efficacy data suggest that the D M P C / C h o l formulation was therapeutically more active in the L S I 8 0 tumour model and was as efficacious as the D S P C / C h o l formulation when treating A431 tumours. These data place emphasis on the importance o f designing liposomal formulations that optimize drug biological availability rather than drug delivery. In Chapter 5, the role of liposomal drug delivery in the treatment of liver localized cancer was investigated. The therapeutic activity of liposomal formulations of vincristine, doxorubicin and mitoxantrone were tested in a model where L1210 tumour cells seed in the liver and the spleen. Only treatment with D M P C / C h o l mitoxantrone at a 10 mg/kg drug dose effected cures as measured by survival beyond 60 days. In order to better understand the activity of mitoxantrone in the liver, the role of drug delivery was assessed. This was modulated through the use of procedures that cause reductions of liposome accumulation in the liver and it was predicted that this would result in decreased therapeutic activity. Reduction in liver accumulation was achieved by either the use of P E G modified lipids or by methods designed to suppress phagocytic cell activity in the liver, referred to as hepatic M P S blockade. Decreases in anti-tumour activity were observed with the P E G formulation; however, the use of M P S blockade failed to reduce the therapeutic activity of D M P C / C h o l mitoxantrone, despite lower drug delivery. These data 137 demonstrate that although the Kupffer cells play a role in liposome accumulation, this population is not responsible for mediating therapeutic activity of D M P C / C h o l mitoxantrone. 6.2 Discussion The results from this thesis highlight the importance of drug release. Drug encapsulated within the liposome is not biologically available and therefore, does not play a role in the therapeutic activity. Formulations which have focused on drug retention have decreased drug toxicity (Mayer et al, 1989) and improved the therapeutic activity (Boman et al, 1994), whereas other formulations exhibiting rapid drug release in the circulation tend to exhibit increased toxicity and a reduction in the therapeutic activity (Mayer et al, 1994). A s demonstrated in Chapter 3, formulations which retain the drug (such as D S P C / C h o l mitoxantrone) can be less effective than the free drug, a consequence of reduced drug biological availability. Hence, liposomal formulations must be optimized in terms of the rate of drug release. It is important to note that stability in the circulation is also a crucial parameter, as drug that is released in the circulation is believed to have a negligible role in the therapeutic activity. Once the liposomes extravasate into the site of tumour development, drug release is required to optimize exposure to the drug. This is in contrast to mechanisms postulated on the basis o f slow release of drug from liposomes that reside in the blood compartment. If drug release within the circulation were a crucial parameter, administration of drug via infusion pumps should yield greater increases in therapeutic activity. Often drug infusion procedures result in only marginal improvements in clinical response (Jackson et al, 1989; 1985). It is, of course, easy to stress potential advantages of liposomal delivery systems because the experience with these formulations is far less when compared to studies in humans that have evaluated infusion 138 approaches. However, the use of liposomes as drug carriers provides a drug reservoir at the site o f tumour development and, i f appropriately designed, these systems w i l l decrease systemic exposure o f the associated anti-cancer drug. Triggered release of drug ideally would occur using liposome systems which retain the drug in the circulation but once the liposome extravasates in the disease site, drug release is stimulated by either an external signal or a change in the liposome. These systems would theoretically improve the therapeutic activity but also decrease toxicity since drug release is emphasized at the region where drug is required. The concept of triggered release has been studied through the use of p H sensitive liposomes which exploit the tumour's acidic interior to cause the liposomes to release their contents at the target site (Aicher et al, 1994; Connor et al., 1984). Thermosensitive liposomes are also being developed where liposomes are injected and regional hyperthermia causes release of the liposome contents (Kakinuma et al, 1996; Gaber et al, 1996; Chelvi et al, 1995). In addition, it has been demonstrated that lipids such as unsaturated P E ' s (which do not normally adopt a bilayer structure) can be stabilized into a bilayer conformation through the use of lipids such as P E G - P E (Holland et al, 1996a). A s the P E G moiety leaves the liposome, the liposome destabilizes, releasing the drug or fuses with the tumour cell (See Figure 6.1). It has been established that P E G modified lipids can be designed to exchange out o f the liposomal membrane (Holland et al, 1996b) or alternatively, the P E G moiety can be lost due to chemical degradation of the lipid (Kirpotin et al, 1996b; Parr et al, 1994). A s demonstrated in Chapter 3 of this thesis, the D M P C / C h o l formulation of mitoxantrone is active in liver localized disease. As noted in Chapter 5, reduction of the therapeutic activity can be attained through the use of PEG-modified lipids, which decreases the accumulation of D M P C / C h o l mitoxantrone in the liver. However, reduction achieved through the use o f M P S 139 Figure 6.1 Future design of liposomes Liposomes are designed to be stable in the circulation. Incorporation of PEG increases circulation lifetime leading to a greater potential for extravasation into the tumor site. Loss of PEG reveals targeting ligand. Increases binding to the target cell. Further loss of PEG causes destabilization of the liposome. Promotes fusion with the target cell and release of drug at the site. 140 blockade, did not result in the same effect. It is surprising to find that the Kupffer cells do not play a role in mediating the activity of liposomal mitoxantrone. It has been demonstrated that the Kupffer cells can act as a reservoir for drugs, releasing the free form back into circulation (Storm et al, 1988). In addition, the use of liposomal doxorubicin and vincristine on the L1210 tumour model demonstrated disappointing results. This was thought to be due to the effects of these drugs on the Kupffer cell population thereby reducing drug delivery to the liver. It can be concluded, however, that the Kupffer cells do not play a role in mediating the activity of liposomal mitoxantrone. The difference between the two methods employed to reduce drug delivery to the liver is specificity. The use of PEG-modified lipids reduces delivery of liposomes to cells due to the steric shielding which inhibits liposome-cell interactions (Du et al, 1997). It is plausible that the use of P E G has altered the distribution of liposomes within the liver, and thus decreasing delivery to the cell population mediating activity. It has been observed that although P E G liposomes can cause an increase in the circulation levels of liposomes, this does not translate to an increase in tumour accumulation (Parr et al, 1997). Similarly in the liver, as the liposome percolate throughout the liver, there are several cell populations which can interact with the liposomes. The use of P E G can inhibit the interactions with these cells; thereby decreasing the therapeutic activity of liposomal mitoxantrone. The use of M P S blockade eliminates only the Kupffer cell population and thus, liposome interaction with this population. Although the use of M P S has excluded the possibility of Kupffer cells, there are still other cell populations which liposomes associate with and this may mediate activity of liposomal mitoxantrone in the liver. For example, the endothelial cells are also capable of phagocytosis and also play a role in liposomal transport. The use of M P S blockade would not affect this transportation role; 141 however, the use o f P E G lipids inhibits interactions of the liposomes with the endothelial cells, thereby reducing transport. A n assessment of delivery to the endothelial cells would be instrumental in determining the involvement of these cells. It has been demonstrated that the Kupffer cells can affect the phagocytic capability of the endothelial cells in the liver (Deaciuc et al, 1994). Thus, the use of M P S blockade may increase the liposome accumulation in the endothelial cell population and in tern, mediate the therapeutic activity of liposomal mitoxantrone. The M P S affects only the Kupffer cell population and there are still circulating pools of monocytes which are still capable of liposome uptake. Future experiments could examine the monocyte population further by reducing the monocyte population through the use o f anti-CD 14 antibody and carbonyl iron (Holtrop et al, 1992). This would determine the role of the M P S rather than focusing on the Kupffer cell population. In addition, the mechanisms by which the liver process liposomes are still under investigation. Three pathways have been proposed by Scherphof et al. (1998), suggesting that liposomal elimination is a very complicated process. Two o f the pathways involve receptor binding (apoE-mediated receptor and an unknown receptor) and then endocytic internalization into the lysosomal compartment. The third involves an H D L receptor and results in transference of certain bilayer constituents to the bilayer o f the hepatocyte. A more thorough examination of how hepatocytes process liposomal mitoxantrone may also explain the therapeutic activity. A s seen in Chapter 5, delivery to the hepatocytes was unaffected by M P S blockade or the use of P E G . Although the P E G inhibits cell interactions, the P E G moiety does not provide complete protection and protein binding w i l l eventually overcome these benefits of surface stabilization. Cells that interact with these liposomes may internalize them, however, the remaining P E G may alter the processing of the liposome. 142 Targeting the liver can also be achieved through the use of charged lipid such as phophatidylserine. The majority of phosphatidylserine containing liposomes accumulate in the liver, and therefore, this would enhance delivery to the site of tumour development. However, the majority of these liposome accumulate in the Kupffer cells (Spanjer et al., 1986) and these cells do not play a role in mediating liposomal mitoxantrone. Thus, it w i l l be interesting to note i f the increased accumulation of the liposomes in the liver w i l l result in the same therapeutic activity i f the majority of the liposomes are taken up by the Kupffer cells. In conclusion, drug release is an important parameter when designing liposomal formulations regardless i f the liposomes accumulate rapidly at the site of tumour development, such as the liver, or at an extravascular site where liposomes accumulate slowly. In addition, delivery to the cells mediating activity is also critical. Although the Kupffer cells are responsible for liposome uptake in the liver, they do not affect the therapeutic activity of D M C P / C h o l mitoxantrone. In closing, the use of liposomal mitoxantrone for the treatment of liver cancer holds much promise and continued studies in the liver's role of processing these carriers could improve liposomal treatment for this disease. 143 REFERENCES Abra, R . M . and Hunt, C A . Liposome disposition in vivo. 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