Rational design of nanocarriers for oral peptide administration

Universidad de Santiago de Compostela Facultad de Farmacia Departamento de Farmacia y Tecnología Farmacéutica Doctoral thesis Rational design of nanocarriers for oral peptide administration (Reduced Version) Zhigao Niu Santiago de Compostela, 2016 Dr. María José Alonso Fernández, Full Professor at the Department of Pharmaceutical Technology in University of Santiago de Compostela Dr. Manuel Jesús Santander Ortega, Assistant Professor at the Department of Medical Sciences in University of Castilla-La Mancha Report: That the experimental dissertation entitled: “Rational design of nanocarriers for oral peptide administration” presented by Zhigao Niu was conducted under their supervision at the Department of Pharmaceutical Technology at the University of Santiago de Compostela. Being completed, they authorize its presentation and evaluation by the assigned jury members. And for the record, they issue and sign the present certificate in Santiago de Compostela, March 10th, 2016. Prof. María José Alonso Fernández Dr. Manuel Jesús Santander Ortega Learning without thought means labor lost. Thought without learning is perilous. Confucius (551 – 479 BC) Acknowledgement In Chinese we have a proverb called “饮水思源 (yin shui si yuan)”, which literally means when one drinks water, one should not forget where it comes from. At the end of this PhD thesis, I would like to express my appreciation to all the people that instructed me and helped me to get through this 4-year-fight. First of all, I would like to express my sincere gratitude to my thesis directors; Prof. María José Alonso and Dr. Manuel Santander Ortega, who opened my eyes to see the fascinating perspective in this nano-pharmaceutical world, led me to delve into the wealth of knowledge and advance to be a meticulous researcher. On the other hand I am very grateful to their responsible revision of this manuscript. Their dedication and passion for the work during this process steered me to realize my potential. My special appreciation goes to Dr. Inmaculada Conejos (USC), Brendan Griffin and Prof. Caitriona O´Driscoll (UCC, Ireland), who made a great contribution to the preparation of thesis chapter 1. At the same time, I would like to thank our research collaborators, who carried out numerous experiments and enriched the thesis chapters 2 and 3: Dr. Federico Benetti (Veneto Nanotech, Italy), Aloïse Mabondzo (CEA, France), Ana Beloqui and Prof. Véronique Préat (UCL, Belgium) for the cell studies; Dr. Patrik Lundquist, Prof. Per Artursson (Uppsala University, Sweden) for the studies on human intestinal tissue, Josep Garcia, Dr. Meritxell Teixidó and Prof. Ernest Giralt (IRB, Spain) for the polymer synthesis, Dr. Ilaria Marigo (IOV, Italy), Sulay Tovar and Prof. Carlos Dieguez (USC) for the animal studies. All the other Trans-Int partners are also much appreciated for broadening my horizon, and the collaboration with Sanofi offered me an opportunity to learn from a world-class pharmaceutical company, which helped me to establish my knowledge of the industrial world. I would like to thank all the members of Nanobiofar family, especially my Trans-Int team: Lungile, Irene, Tamara, Matilde, as well as Inma, Ana and Adam, who used to work with us. Just like we said in the video: “We are a team. We have a goal. We are Trans-Int.” The light into the night in CIMUS witnessed the eventful years we came through together. I will always be with you, today, tomorrow, and forever. Behind all our achievements, there was the direction from Prof. María José Alonso, Dolores Torres, Dr. Noemi Csaba and Desirée Teijeiro. Also I have to mention Dr. Marcos Garcia Fuentes, María de la Fuente and María Victoria Lozano, whom I used to ask for advice. I am also grateful to my other lab-mates that cared for me, helped me and encouraged me tirelessly: Sonia, Adriana, Carmen, Ana Cadete, José, Belén Lopez, Hsu Wei-Hsin (Howl), Eleni, Sara Cordeiro, Sara Vicente, Jorge, José Vicente, Giovanna, Raquel, Vanesa, Ana Olivera, Carla, Andrea, Sofia, Nataliya, Fernando, Khair, Blanca, Marta, Surasa, Mariajo, Maruthi, Ivana, and so forth. Of course, I will not forget the technical supports from Belén Cuesta, Rafa, David Gonzales (animal studies), and the administrative help from Puri. It is complicated to work in a foreign country, but all their professional aid and kindness facilitated my research here. Last but not the least, I would also like to thank Prof. Jean-Pierre Benoit and Frank Boury, who gave me the chance to spend a couple of months in Inserm U1066, (Angers), where I learnt new techniques and met new friends. I appreciate the supervision of Prof. Frédéric Lagarce, Jean-Christophe Gimel and the help from Dr. Carl Simonsson, Hélène Malhaire, Giovanna Lollo, Emilie André, Sai Krishna, and Clément Haeck during my stay there in France. Thank you to all the people I met in Santiago de Compostela, this rainy but lovely city. Your company really made me feel at home. One day if I leave Galicia, I will have “morriña” as you do, for sure. Indice Resumen - Abstract .................................................................................... 3 Resumen ............................................................................................... 5 Abstract ................................................................................................ 7 Introducción................................................................................................ 9 Introduction .............................................................................................. 19 Chapter 1 - Lipid-Based Nanocarriers for Oral Peptide Delivery ........... 29 Antecedentes, Hipótesis y Objetivos........................................................ 97 Background, Hypothesis and Objectives ................................................. 97 Chapter 2 - Polyarginine Nanocapsules: a Potential Carrier for Oral Peptide Delivery ..................... 109 Chapter 3 - Polymer Nanocomplexes for Oral Insulin Delivery ........... 149 Overall discussion .................................................................................. 155 Conclusiones - Conclusions ................................................................... 169 Resumen - Abstract 8 studies showed the capacity of the nanocomplexes to remarkably enhance insulin transport (47.59% cell uptake and 2.11% transport). Overall, the results of these thesis underline the possibility to engineer nanocarriers that are able to overcome the multiple biological barriers associated to the oral peptide administration. Introducción 9 Introducción Introducción 10 Introducción 11 Introducción El reto asociado a la administración oral de péptidos En las últimas décadas se ha visto claramente cómo, en general, la terapia basada en el uso de péptidos y proteínas podría verse claramente beneficiada mediante el diseño de plataformas que posibiliten su administración por vía oral. Sin embargo, las barreras biológicas que los péptidos han de superar para lograr su absorción sistémica tras su administración por vía oral suponen un gran reto tecnológico. Entre estas moleculas se encuentra la insulina, probablemente uno de los fármacos más complejos en cuanto a su formulación en formas de administración oral (BCS III). Descubierta en el año 1921 (Banting y Best), el uso terapéutico de la insulina se vio particularmente incrementado al lograr suproducción masiva mediante técnicas de ADN recombinante (1). La posibilidad de administrar insulina por vía oral es particularmente atractiva, ya que además de una mayor aceptación por parte del paciente, la absorción de insulina a través del intestino podría imitar la circulación enterohepática de insulina endógena, haciendo posible un efecto sostenido del fármaco. En conjunto, estas ventajas representarían un gran beneficio para el tratamiento de una enfermedad crónica como la diabetes. No obstante, esta posibilidad está aún lejos de ser alcanzada debido a las diferentes barreras biológicas que presenta el tracto gastrointestinal (GIT). Estas barreras incluyen los fluidos intestinales ricos en enzimas, que pueden degradar al péptido, la capa de moco que protege el epitelio intestinal y el propio epitelio, que no permite el paso de macromoléculas hidrófilas (2-4). Mientras que el problema de la estabilidad podría abordarse mediante el uso de recubrimientos entéricos, la baja permeabilidad de la mucosa intestinal sigue siendo la limitación de las formulaciones desarrolladas hasta el momento para su paso a clínica (5). Las diferentes estrategias utilizadas para resolver estos problemas incluyen la modificación química del péptido, la co-administración de promotores de la absorción e inhibidores de enzimas, Introducción 12 encapsulación del péptido en sistemas de liberación adecuados (3). Desafortunadamente, se ha visto que hasta ahora estas estrategias poseen ciertas limitaciones que podrían justificar su limitado desarrollo clínico. Por ejemplo, la modificación química de péptidos puede inducir cambios en el perfil de actividad/toxicidad de la molécula original (6). La co-administración de insulina con inhibidores enzimáticos y promotores de la absorción puede producir efectos secundarios no deseados tales como una digestión lenta, o la absorción intestinal de compuestos no deseados. El desarrollo de nuevos nanotransportadores capaces de ayudar a los péptidos a superar las barreras indicadas supone un objetivo atractivo, no carente de desafíos técnicos significativos. Nanovehículos y sistemas lipídicos para administración oral de fármacos A día de hoy la nanotecnología ha permitido la administración oral de fármacos tanto solubles como insolubles, así como la administración dirigida a células intestinales específicas y una mejor absorción intestinal del fármaco a través de las vías paray transcelular (7). Diferentes nanotransportadores, entre los que se incluyen nanopartículas poliméricas y lipídicas, liposomas, micro / nanoemulsiones, sistemas auto-emulsionables micro / y nanocápsulas han sido diseñados para la administración oral de insulina. La mayoría de los cuales se han tratado en el Capítulo 1 de esta tesis, titulado "Lipid-based nanocarriers for oral peptide delivery". Haciendo uso de las propiedades específicas de los polímeros, lípidos y otros excipientes auxiliares, ha sido posible diseñar nanovehículos que exhiben una buena estabilidad coloidal en el entorno GIT, además de poseer la capacidad de proteger a la insulina frente al ataque enzimático, así como de controlar su liberación y favorecer suabsorción a través del epitelio intestinal. Sin embargo, hay muy pocos estudios centrados en el estudio de la interacción de estos sistemas con las barreras biológicas asociadas a la vía oral de una manera sistémica. Es por esto, quizás, que existe una variabilidad significativa y una reproducibilidad limitada en los resultados in vivo publicados hasta el momento. En general, este escenario hace evidente la necesidad de llevar a cabo estudios Introducción 13 mecanísticos que nos ayuden a entender el mecanismo de interacción entre los nanotransportadores de fármacos y las barreras biológicas asociadas a la vía oral. Entre los enfoques tecnológicos explorados para lograr la administración de insulina por vía oral, los que emplean excipientes lipídicos son particularmente prometedores debido a su gran diversidad, biocompatibilidad y una funcionalidad específica. La mayoría de los excipientes lipídicos son derivados de aceites o grasas consumidos normalmente en la dieta, que confieren, además de la biodegradabilidad y por tanto la baja toxicidad, la capacidad de (i) aumentar la permeabilidad de la membrana intestinal, (ii) reducir la degradación proteolítica y (iii) aumentar transporte linfático intestinal. En este campo, diversos nanotransportadores lipídicos han sido estudiados para la administración oral de péptidos, incluyendo nanopartículas lipídicas sólidas, liposomas, microemulsiones, nanoemulsiones, sistemas de administración de fármacos autoemulsionantes y las micro / nanoemulsiones recubiertas por un polímero, es decir, las nanocápsulas. Debido a su naturaleza hidrofóbica un reto a superar ha sido la baja eficiencia de encapsulación de la insulina en estos nanosistemas. Varias estrategias se han explorado para solucionar esta limitación. Una opción muy común ha sido el uso de ácidos grasos de cadena media y sus glicéridos, que se dispersan fácil en medios acuosos y son capaces de solubilizar péptidos (8, 9). Otras estrategias de formulación incluyen (i) la formación de emulsiones W/O/W (10, 11), (ii) la hidrofobización de los péptidos para promover la formación de micelas inversas (12, 13) o la complejación / conjugación con restos lipófilos (14) y (iii) la interacción electrostática con los lípidos o agentes tensioactivos (10). Todas estas alternativas han llevado a una mejora significativa de la capacidad de encapsulación de los sistemas lipídicos. Los mecanismos involucrados en la liberación del péptido desde la matriz lipídica incluyen (i) la difusión simple a través de los canales que contiene la matriz de lípidos (15), (ii) la disociación iónica seguida por una difusión a través de los canales (16, 17), (iii) la degradación de la matriz lipídica mediada por lipasas (15), y (iv) Hinchado-rotura de los nanosistemas por fuerzas osmoticas (18). A parte de la Introducción 14 liberación controlada de fármacos, los lípidos generalmente tienen una alta afinidad por la membrana celular, y muchos de los ingredientes lipídicos de estos nanosistemas han mostrado propiedades como promotores de la absorción, i.e. los triglicéridos de C10 y C8 (19, 20). Finalmente, estos sistemas podrían mejorar su comportamiento in vivo mediante el uso de una cubierta polimérica (dando lugar a la formación de nanocápsulas) que mejorase la estabilidad de los nanotransportadores en condiciones fisiológicos. Poliarginina Los polipéptidos y poliaminoacidos ricos en arginina, tales la poliarginina, han sido ampliamente estudiados en el desarrollo de nanovehículos para la administración oral de insulina debido a sus excelentes propiedades como promotores de la absorción celular. Estudios mecanísticos en los que se compararon moléculas ricas en arginina con moléculas ricas en lisina, mostraron que las primeras lograban una mayor eficiencia de internalización, lo cual indica claramente la importancia de los grupos guanidinio de la cadenas laterales de poli/oligoarginina (21). Además de la alta ionización de la cadena lateral de guanidina (arginina pKa ~ 12.5) (22), la mayor eficiencia se atribuye al hecho de que los grupos ricos en guanidinio son capaces de formar puentes de hidrógeno en la superficie celular. La formación de estos enlaces convierte al péptido hidrófilo en un material más hidrófobo, reduciendo la energía necesaria para pasar a través de la bicapa lipídica de las células (23). La poli-L-arginina (PARG), polipéptido que cuenta con un máximo de 554 residuos de arginina. El potencial de esta molécula para mejorar la absorción de fármacos hidrófilos se ha estudiado tanto a nivel de mucosas como a nivel intracelular (24-30). Estudios llevados a cabo en ratas han puesto de relación el efecto del peso molecular de la PARG en la absorción FD-4 (molécula modelo). En función de su capacidad para mejorar la absorción de FD-4 las diferentes PARG se ordenaron de la siguiente manera: PARG 92 kDa> PARG 45,5kDa> PARG 8,9 kDa (29). Interesante es también Introducción 15 el hecho de que todas las PARG estudiadas mostraron una mayor capacidad para promover la absorción que promotores clásicos como son el glicolato sódico o el taurocolato sódico (29). En esta línea, la PARG ha mostrado un gran potencial para mejorar la absorción de péptidos y proteínas (con un peso molecular igual o inferior a 20KDa) a través de mucosas (30). Conclusiones and perspectivas La PARG ha mostrado un gran potencial para mejorar el transporte de péptidos a través de las células epiteliales, ya sea a través de la vía endocítica o por una ruta paracelular. También se han realizado estudios centrados en el uso de nanotransportadores que contenían PARG para mejorar el transporte intracelular del fármaco encapsulado. No obstante, hay pocos estudios relacionados con su uso para la administración oral de péptidos. En general, los estudios publicados en los últimos años señalan la posibilidad de combinar de forma sinérgica la nanotecnología con el uso de polímeros ricos en arginina, con el objeto de mejorar el transporte de macromoléculas a través de la barrera intestinal. Ésta ha sido la base principal del trabajo desarrollado en esta tesis. Esta idea se ha materializado mediante la combinación de macromoléculas ricas en arginina con otros promotores de la absorción (sales biliares, aceites), en forma de nanoestructuras capaces de encapsular insulina y mejorar su biodisponibilidad tras su administración oral. Esta breve introducción se complementa con la revisión presentada en esta tesis (Capítulo 1) titulada "Lipid-based nanocarriers for oral peptide delivery", recientemente enviada para su publicación en la revista Advanced Drug Delivery Reviews. Introducción 16 Referencias 1. Brange J, Vølund A. Insulin analogs with improved pharmacokinetic profiles. Advanced Drug Delivery Reviews. 1999;35(2–3):307-35. 2. Wang J, Yadav V, Smart AL, Tajiri S, Basit AW. Toward oral delivery of biopharmaceuticals: an assessment of the gastrointestinal stability of 17 peptide drugs. Mol Pharm. 2015;12(3):966-73. 3. Herrero EP, Alonso MJ, Csaba N. Polymer-based oral peptide nanomedicines. Therapeutic Delivery. 2012;3(5):657-68. 4. Morishita M, Peppas NA. Is the oral route possible for peptide and protein drug delivery? Drug Discov Today. 2006;11(19-20):905-10. 5. Al-Hilal TA, Alam F, Byun Y. Oral drug delivery systems using chemical conjugates or physical complexes. Adv Drug Deliv Rev. 2013;65(6):845-64. 6. Duncan R, Vicent MJ. Polymer therapeutics-prospects for 21st century: the end of the beginning. Adv Drug Deliv Rev. 2013;65(1):60-70. 7. Plapied L, Duhem N, des Rieux A, Préat V. Fate of polymeric nanocarriers for oral drug delivery. Current Opinion in Colloid & Interface Science. 2011;16(3):228-37. 8. Gershanik T, Benita S. Self-dispersing lipid formulations for improving oral absorption of lipophilic drugs. European Journal of Pharmaceutics and Biopharmaceutics. 2000;50(1):179-88. 9. Griffin B, O'Driscoll C. Opportunities and challenges for oral delivery of hydrophobic versus hydrophilic peptide and protein-like drugs using lipid-based technologies. Therapeutic Delivery. 2011;2(12):1633–53. 10. Garcia-Fuentes M, Torres D, Alonso MJ. Design of lipid nanoparticles for oral delivery of hydrophilic macromolecules. 2003;27(2-3):159-68. 11. Anton N, Benoit JP, Saulnier P. Design and production of nanoparticles formulated from nano-emulsion templates-a review. J Control Release. 2008;128(3):185-99. 12. Chen C, Fan T, Jin Y, Zhou Z, Yang Y, Zhu X, et al. Orally delivered salmon calcitonin-loaded solid lipid nanoparticles prepared by micelle–double emulsion method via the combined use of different solid lipids. Nanomedicine. 2012;8(7):1085-100. 13. Liu J, Gong T, Wang C, Zhong Z, Zhang Z. Solid lipid nanoparticles loaded with insulin by Introducción 17 sodium cholate-phosphatidylcholine-based mixed micelles: preparation and characterization. Int J Pharm. 2007;340(1-2):153-62. 14. Yuan H, Jiang SP, Du YZ, Miao J, Zhang XG, Hu FQ. Strategic approaches for improving entrapment of hydrophilic peptide drugs by lipid nanoparticles. Colloids Surf B Biointerfaces. 2009;70(2):248-53. 15. Christophersen PC, Zhang L, Yang M, Nielsen HM, Mullertz A, Mu H. Solid lipid particles for oral delivery of peptide and protein drugs I--elucidating the release mechanism of lysozyme during lipolysis. Eur J Pharm Biopharm. 2013;85(3 Pt A):473-80. 16. Garcia-Fuentes M, Torres D, Alonso MJ. New surface-modified lipid nanoparticles as delivery vehicles for salmon calcitonin. Int J Pharm. 2005;296(1-2):122-32. 17. Oh KS, Han SK, Lee HS, Koo HM, Kim RS, Lee KE, et al. Core/Shell Nanoparticles with Lecithin Lipid Cores for Protein Delivery. Biomacromolecules. 2006;7(8):2362-7. 18. Cunha AS, Grossior JL, Puisieux F, Seiller M. Insulin in w/o/w multiple emulsions: Preparation, characterization and determination of stability towards proteases in vitro. Journal of Microencapsulation. 1997;14(3):311-9. 19. Brayden DJ, Gleeson J, Walsh EG. A head-to-head multi-parametric high content analysis of a series of medium chain fatty acid intestinal permeation enhancers in Caco-2 cells. European Journal of Pharmaceutics and Biopharmaceutics. 2014;88(3):830-9. 20. Brayden DJ, Walsh E. Efficacious Intestinal Permeation Enhancement Induced by the Sodium Salt of 10-undecylenic Acid, A Medium Chain Fatty Acid Derivative. The AAPS Journal. 2014;16(5):1064-76. 21. Mitchell DJ, Steinman L, Kim DT, Fathman CG, Rothbard JB. Polyarginine enters cells more efficiently than other polycationic homopolymers. The Journal of Peptide Research. 2000;56(5):318-25. 22. Nakase I, Takeuchi T, Futaki S. Cell Penetrating Peptides for Chemical Biological Studies. In: Langel Ü, editor. Cell-Penetrating Peptides: Methods and Protocols. New York, NY: Springer New York; 2015. p. 387-96. 23. Rothbard JB, Jessop TC, Lewis RS, Murray BA, Wender PA. Role of Membrane Potential and Hydrogen Bonding in the Mechanism of Translocation of Guanidinium-Rich Peptides into Cells. Introduction 24 across the cell membrane and even across mucosal tissues (16-20). A comparative study in vivo aimed at determining the importance of the molecular weight (Mw) of PARG on its permeation enhancer capacity following nasal delivery (FD-4 as model drug) resulted in the following permeability ranking: PARG 92 kDa > PARG 45.5 kDa > PARG 8.9 kDa (21). Moreover, these authors showed that the permeation enhancing capacity of PARG polymers, was higher than that of classical permeation enhancers such as sodium glycocholate or sodium taurocholate (21). This capacity is in agreement with the enhancement of the transport of various protein drugs across nasal membrane (22). Additionally, it has been reported that the PARG penetration enhancing effect is dose-dependent (23). PARG has a good safety profile, without causing damage to rat erythrocyte and isolated rabbit nasal mucosa (24). However, opposite to the ability displayed by most CPP to enhance endocytosis, PARG polymers are rather known to induce the transient opening of cellular tight junctions, thereby favoring the paracellular drug absorption in a molecular weight dependent fashion (24). As later revealed by mechanistic studies, these polymers can lead to transient internalization of tight junction proteins between the epithelium cells via clathrin-mediated endocytosis, which leads to the increase of the permeability of hydrophilic drugs (25-27). Finally, the internalized tight junction proteins are then recycled through endosome pathway, re-forming the intercellular tight junction. These features are similar to the paracellular permeation enhancing properties displayed by chitosan, which has been widely employed for pharmaceutical applications including drug delivery systems and turned out to be a great success (28). PARG has been compared with regard to their capacity to interact with the anionic phospholipids of liposomes. The results showed that as the molecular weight rises from 8 up to 550 arginine residues, the interaction with the cell membrane also increases. Incubation of PARG polymers, with around 300 arginine residues or more, Introduction 25 with lipid membranes resulted in the lipid membrane perturbation and the subsequent translocation across the phospholipid bilayers (29). These results are, somehow, in contradiction with those claiming that the ideal number of arginine residues is in the range of 7-15. Therefore, overall, the limited amount of literature regarding the use of arginine-rich polymers/polymers in the design of antigen delivery carriers suggest the interest to further investigate the potential of these promising biomaterials. Conclusions and perspectives PARG has exhibited a capacity to transport peptides across epithelial cells either through an endocytic or paracellular routes. Moreover, there are a few articles disclosing preliminary data about the potential of PLGA and lipid nanoparticles involving arginine for enhancing insulin absorption (30, 31). On the other hand, PARG-based nanocarriers have also been designed for intracellular drug delivery and few attempts have been directed towards their use for oral peptide delivery. Overall, the background literature suggest that there are possibilities to synergically combine nanotechnology with the use of arginine-rich polymers with the final goal of improving the transport of macromolecules across the intestinal barrier (32), and this has been the main basis for the idea behind this thesis work. This idea was attempted to be materialized by combining polyarginine with other penetration enhancers (bile salts, oils) as well as protective PEGylated polymers. This brief introduction is complementary to the review presented in thesis chapter 1 entitle “Lipid-based nanocarriers for oral peptide delivery”, recently submitted for publication in the Advanced Drug Delivery Reviews. Introduction 26 References 1. Brange J, Vølund A. Insulin analogs with improved pharmacokinetic profiles. Advanced Drug Delivery Reviews. 1999;35(2–3):307-35. 2. Mitragotri S, Burke PA, Langer R. Overcoming the challenges in administering biopharmaceuticals: formulation and delivery strategies. Nat Rev Drug Discov. 2014;13(9):655-72. 3. Renner DB, Svitak AL, Gallus NJ, Ericson ME, Frey WH, Hanson LR. Intranasal delivery of insulin via the olfactory nerve pathway. Journal of Pharmacy and Pharmacology. 2012;64(12):1709-14. 4. Fonte P, Araújo F, Reis S, Sarmento B. Oral Insulin Delivery: How Far are We? Journal of Diabetes Science and Technology. 2013;7(2):520-31. 5. Kim ES, Plosker GL. AFREZZA® (insulin human) Inhalation Powder: A Review in Diabetes Mellitus. Drugs. 2015;75(14):1679-86. 6. Wang J, Yadav V, Smart AL, Tajiri S, Basit AW. Toward oral delivery of biopharmaceuticals: an assessment of the gastrointestinal stability of 17 peptide drugs. Mol Pharm. 2015;12(3):966-73. 7. Herrero EP, Alonso MJ, Csaba N. Polymer-based oral peptide nanomedicines. Therapeutic Delivery. 2012;3(5):657-68. 8. Morishita M, Peppas NA. Is the oral route possible for peptide and protein drug delivery? Drug Discov Today. 2006;11(19-20):905-10. 9. Al-Hilal TA, Alam F, Byun Y. Oral drug delivery systems using chemical conjugates or physical complexes. Adv Drug Deliv Rev. 2013;65(6):845-64. 10. Duncan R, Vicent MJ. Polymer therapeutics-prospects for 21st century: the end of the beginning. Adv Drug Deliv Rev. 2013;65(1):60-70. 11. Plapied L, Duhem N, des Rieux A, Préat V. Fate of polymeric nanocarriers for oral drug delivery. Current Opinion in Colloid & Interface Science. 2011;16(3):228-37. 12. Gonzalez-Aramundiz JV, Lozano MV, Sousa-Herves A, Fernandez-Megia E, Csaba N. Polypeptides and polyaminoacids in drug delivery. Expert Opin Drug Deliv. 2012;9(2):183-201. 13. Mitchell DJ, Steinman L, Kim DT, Fathman CG, Rothbard JB. Polyarginine enters cells more efficiently than other polycationic homopolymers. The Journal of Peptide Research. 2000;56(5):318-25. Introduction 27 14. Nakase I, Takeuchi T, Futaki S. Cell Penetrating Peptides for Chemical Biological Studies. In: Langel Ü, editor. Cell-Penetrating Peptides: Methods and Protocols. New York, NY: Springer New York; 2015. p. 387-96. 15. Rothbard JB, Jessop TC, Lewis RS, Murray BA, Wender PA. Role of Membrane Potential and Hydrogen Bonding in the Mechanism of Translocation of Guanidinium-Rich Peptides into Cells. Journal of the American Chemical Society. 2004;126(31):9506-7. 16. Lozano MV, Lollo G, Alonso-Nocelo M, Brea J, Vidal A, Torres D, et al. Polyarginine nanocapsules: a new platform for intracellular drug delivery. Journal of Nanoparticle Research. 2013;15(3). 17. Harashima H, Futaki S, Kogure K. Polyarginine-modified liposome having nuclear entry ability. Google Patents; 2013. 18. EP H, VZ S, DT K, TJ D. Polyarginine segments in block copolypeptides drive both vesicular assembly and intracellular delivery. Natura Materials. 2007;6(1):52-7. 19. Rawat A, Yang T, Hussain A, Ahsan F. Complexation of a Poly-l-Arginine with Low Molecular Weight Heparin Enhances Pulmonary Absorption of the Drug. Pharmaceutical Research. 2007;25(4):936-48. 20. Ong J, Jennings R, Stetsko G. Methods and compositions for enhanced transmucosal delivery of peptides and proteins. Google Patents; 2011. 21. Natsume H, Iwata S, Ohtake K, Miyamoto M, Yamaguchi M, Hosoya K-i, et al. Screening of cationic compounds as an absorption enhancer for nasal drug delivery. International Journal of Pharmaceutics. 1999;185(1):1-12. 22. Miyamoto M, Natsume H, Satoh I, Ohtake K, Yamaguchi M, Kobayashi D, et al. Effect of poly-l-arginine on the nasal absorption of FITC-dextran of different molecular weights and recombinant human granulocyte colony-stimulating factor (rhG-CSF) in rats. International Journal of Pharmaceutics. 2001;226(1–2):127-38. 23. Miyamoto M, Natsume H, Iwata S, Ohtake K, Yamaguchi M, Kobayashi D, et al. Improved nasal absorption of drugs using poly-l-arginine: effects of concentration and molecular weight of poly-l-arginine on the nasal absorption of fluorescein isothiocyanate–dextran in rats. European Journal of Pharmaceutics and Biopharmaceutics. 2001;52(1):21-30. Introduction 28 24. Ohtake K, Natsume H, Ueda H, Morimoto Y. Analysis of transient and reversible effects of poly-l-arginine on the in vivo nasal absorption of FITC-dextran in rats. Journal of Controlled Release. 2002;82(2–3):263-75. 25. Yamaki T, Kamiya Y, Ohtake K, Uchida M, Seki T, Ueda H, et al. A mechanism enhancing macromolecule transport through paracellular spaces induced by Poly-L-Arginine: Poly-L-Arginine induces the internalization of tight junction proteins via clathrin-mediated endocytosis. Pharm Res. 2014;31(9):2287-96. 26. Ohtake K, Maeno T, Ueda H, Ogihara M, Natsume H, Morimoto Y. Poly-L-arginine enhances paracellular permeability via serine/threonine phosphorylation of ZO-1 and tyrosine dephosphorylation of occludin in rabbit nasal epithelium. Pharm Res. 2003;20(11):1838-45. 27. Nemoto E, Takahashi H, Kobayashi D, Ueda H, Morimoto Y. Effects of Poly-L-arginine on the Permeation of Hydrophilic Compounds through Surface Ocular Tissues. Biological and Pharmaceutical Bulletin. 2006;29(1):155-60. 28. Garcia-Fuentes M, Alonso MJ. Chitosan-based drug nanocarriers: where do we stand? J Control Release. 2012;161(2):496-504. 29. Takechi Y, Yoshii H, Tanaka M, Kawakami T, Aimoto S, Saito H. Physicochemical Mechanism for the Enhanced Ability of Lipid Membrane Penetration of Polyarginine. Langmuir. 2011;27(11):7099-107. 30. Liu X, Liu C, Zhang W, Xie C, Wei G, Lu W. Oligoarginine-modified biodegradable nanoparticles improve the intestinal absorption of insulin. International Journal of Pharmaceutics. 2013;448(1):159-67. 31. Tsiourvas D, Sideratou Z, Sterioti N, Papadopoulos A, Nounesis G, Paleos CM. Insulin complexes with PEGylated basic oligopeptides. Journal of Colloid and Interface Science. 2012;384(1):61-72. 32. Salzano G, Torchilin VP. Intracellular Delivery of Nanoparticles with Cell Penetrating Peptides. In: Langel Ü, editor. Cell-Penetrating Peptides: Methods and Protocols. New York, NY: Springer New York; 2015. p. 357-68. Chapter 1 29 Chapter 1 Lipid-Based Nanocarriers for Oral Peptide Delivery This work is done in collaboration with: Inmaculada Conejos-Sánchez 1, Brendan Griffin 2, Caitriona O’Driscoll 2 1. Nanobiofar group, Center for Research in Molecular Medicine and Chronic Diseases, University of Santiago de Compostela, Santiago de Compostela, Spain 2. School of Pharmacy, University College Cork, Cavanagh Pharmacy Building, Cork, Ireland Chapter 1 30 Chapter 1 31 Abstract This chapter is aimed to overview the lipid-based nanostructures designed so far for the oral administration of peptides and proteins, and to analyse the influence of their composition and physicochemical (particle size, zeta potential) and pharmaceutical (drug loading and release) properties, on their interaction with the gastro-intestinal environment, and the subsequent PK/PD profile of the associated drugs. The ultimate goal is to highlight and comparatively analyse the key factors that may be determinant of the success of these nanocarriers for oral peptide delivery. The article ends with some prospects on the challenges to be addressed for the intended commercial success of these delivery vehicles. Abbreviations: BA: bioavailability, BSA: bovine serum albumin, FDA: Food and Drug Administration, GMO: glyceryl monostearate, GRAS: Generally recognized as safe, HLB: hydrophilic-lipophilic balance, LBDDS: lipid-based drug delivery system(s), LC: long chain, LCT: long chain triglycerides, LFCS: lipid formulation classification system, MC: medium chain, MCT: medium chain triglycerides, MW: molecular weight, NC: nanocapsules, OVA: ovalbumin, PACA: poly(alkylcyanoacrylate), p/p: peptides and proteins, PA: pharmacological bioavailability; PD: pharmacodynamics, PEG: poly(ethylene glycol), PK: pharmacokinetics, SA: stearic acid, SC: subcutaneously, sCT: salmon calcitonin, SDC: sodium deoxycholate, SGC: sodium glycocholate, SGF: simulated gastric fluid, STC: sodium taurocholate, SLN: solid lipid nanoparticles, SEDDS: self-emulsifying drug delivery systems, SMEDDS: self-microemulsifying drug delivery systems, SNEDDS: self-nanoemulsifying drug delivery systems, TG: triglycerides, TP: tripalmitin. Chapter 1 32 1. Introduction Since the launch of recombinant human insulin, protein/peptide therapeutics have gained increasing attention as an alternative to conventional small organic drug molecules. With hundreds of peptide/proteins molecules now commercialized, it is estimated that the market value of protein/peptide therapeutics will be €180 billion 2018 (1, 2). Despite this commercial success, it is widely acknowledged that the full clinical potential of these potent macromolecules has been greatly hampered by their necessity of parenteral administration. As a consequence, the search for new strategies, which may enable the oral delivery of peptide/proteins, is one of the main challenges in the drug delivery field. Among the technological approaches explored to achieve efficient oral peptide delivery, those employing lipid excipients are particularly promising because of their wide diversity, favourable biocompatibility and specific functionality. In particular, the fact that the majority of lipids excipients are derived from dietary oils/fats, confers the advantages both in terms of biodegradability and the capacity to cross the intestinal barrier. Starting with an overview of the physicochemical properties of the lipid materials available for drug delivery, this chapter will comparatively analyse the state-of-the-art of nanostructured lipid-based systems intended for oral peptide delivery. The focus will be in four distinct categories of nanocarriers: (i) solid lipid nanoparticles (SLN), (ii) micro and nanoemulsions, including self-emulsifying systems, (iii) liposomes and (iv) hybrid lipid-polymer systems, e.g. nanocapsules. These delivery carriers will be analysed with regard to their capacity to associate and control the release of peptides, and to overcome the multiple biological barriers associated with oral administration. These barriers include (i) the intestinal fluids, which may compromise the stability of the nanocarriers and also that of the associated peptides, (ii) the mucus layer, which may hinder the access of the nanocarriers to the absorbing epithelium and, finally, (iii) Chapter 1 33 the intestinal epithelium. At the end, an attempt will be made to establish a relationship between the nanocarrier biopharmaceutical properties and the effectiveness in terms of PK/PD of the associated peptide 2. Lipids used in the formulation of oral peptide delivery systems Defined by their intrinsic natural origin, lipids constitute a family of molecules largely exploited in the pharmaceutical field due to their favourable physicochemical and biopharmaceutical properties, as described in the following sections. 2.1 Lipids selection A fundamental understanding of the chemistry and physicochemical properties of the lipids is a prerequisite for successful formulation of peptides and proteins. Criteria for the selection of lipids include purity and chemical stability, solvent capacity, water miscibility, digestibility and fate of digested products, safety and regulatory profile (3). Lipids are often “used as received” and due to the diverse array of commercial lipid raw material suppliers, special attention must be given to the datasheet specifications as well as to the batch-to-batch reproducibility (4). Figure 1 and table 1 display the chemical structure and characteristics of lipids commonly used in the formulation of lipid-based nanocarriers intended for oral peptide delivery (5, 6). These include fatty acids (FA), fatty alcohols, long chain (LC) and medium chain (MC) monoglycerides, diglycerides and triglycerides (TG) and phospholipids. Structurally, these amphiphillic molecules possess a hydrophobic region, which is composed of one, two or three hydrocarbon chain(s) of different lengths, and a differentiated polar head. The number and length of the hydrocarbon chain, as well as the degree of unsaturation (double bonds) determine the hydrophobicity of the molecule. The polar nature is associated to the carboxylor alcohol group in the case of fatty acids or fatty alcohols, respectively and to the sterification of the fatty acids with glycerol groups, in the case of the glycerides: Chapter 1 40 peptide drugs (53-56), however, the transformation of these gels into nanometric structures that could work as peptide carriers is still in a premature stage. Table 2. Most common lipids used in the design of lipid-based nanocarriers for oral peptide administration. In the following subsections, the specific details regarding the encapsulation and controlled release properties of lipid nanostructures will be presented. Chapter 1 41 Figure 2. Illustration of some strategies developed to enhance drug loading capacity into lipid based nanocarriers. 3.1.1. Solid Lipid Nanoparticles Solid lipid nanoparticles (SLN) are made of natural, semi-synthetic or synthetic lipids including triglycerides, partial glycerides, fatty acids, waxes, phospholipids and steroids. They additionally incorporate an emulsifier layer, which provides them with the adequate stability upon dispersion in water. SLN are characterized by the fact that they remain in the solid state at room and body temperatures (57). Introduced by Müller et al. and Gasco et al. over 25 years ago (58, 59), the intrinsic lipidic nature of SLN has conferred these systems with a number of interesting properties, namely, (i) Chapter 1 42 the capacity to protect drugs from degradation, owing to structure dependent lipolysis resistance (60, 61), (ii) the possibility to control drug release benefiting from the varied amount and type of lipid employed (61) (ii) a favourable biocompatibility (62, 63); and (v) ease of large scale production (64). As a consequence, a number of formulations, such as TrabiOral™, Rifamsolin™, Ocusolin™, Vansolin™, Zysolin™ are currently under early clinical development (65), however, the advances so far have been mainly focused on hydrophobic drugs, whereas the delivery of peptides and proteins is still at an early phase. Table 3 illustrates examples of SLNs intended for oral peptide delivery. Among the techniques described for the production of SLN, only those that avoid the use of high temperature are, in principle, appropriate for protein/peptide encapsulation. These techniques involve the dissolution of the lipids in organic solvents followed by their diffusion (66-68) and evaporation (69). The formation of w/o/w double emulsions (70) have been the techniques most frequently reported for achieving significant loadings of peptides and proteins within SLN. For example, our group has previously reported the efficient encapsulation (90%) of salmon calcitonin (sCT) by the double emulsion method, profiting from the electrostatic interaction between the cationic peptide, dissolved in the internal aqueous phase, and the anionic lipids, tripalmitin (TP) and lecithin (71). Other authors have observed similar entrapment for sCT (90%) and insulin (AE 98%, drug loading 18.92%) using the w/o/w technique, however, the high loading was attributed to the formation of sodium cholate reverse micellar structures that enveloped the drug (72, 73). The hydrophobic ion pairing (HIP) technique has also been employed to enhance peptide hydrophobicity, and, thus, its entrapment within the lipid matrix. Yuan et al. obtained 75% leuprolide entrapment by complexing with stearic acid (SA) or sodium stearate (SS) prior to the encapsulation within SLN (74). Overall, a number of studies, (Table 3) have achieved good entrapment of peptide drugs making use of the strategies described above. The final loading of the SLN has been found to be dependent on the liposolubility of the Chapter 1 43 peptide (e.g. 33 % loading for levothyroxine (75), but also on the entrapment strategy, having reached loading values in the range of 10-20% for sCT and insulin (72, 73). Chapter 1 44 Peptide Lipids Size and Z-potential Drug entrapment In vitro drug release Ref. Insulin GMO or cetyl palmitate 180nm-1.5μm -7 to -37mV AE up to 84% DL up to 1.23% ~ 20% burst release in 0.01M HCl; slower release in pH 7.4 buffer (66-68 , 76) Cetyl palmitate 350nm AE >43% Biphasic, initial burst and prolonged release (data not shown) (77) Stearic acid, palmitic acid, SPC 110-119nm -49 to -54mV AE 97.8% DL 18.9% 1h release < 30% in PBS; sustained release up to 144h (73) SA-r8, soybean phospholipids 162nm +29.9mV AE 76% DL 3.19% - (78) Stearic acid 58-75nm -13 to -15mV AE 18-40% - (48) Tripalmitin 111-395nm -17 to -39mV - 45% release in pH 7.4 buffer in 1h (70) Hydrogenated castor oil 397-513nm -23 to -30mV AE 78-86% DL 1.58-1.71% Incorporation of PLGA into the SLN offered controlled drug release (79) sCT Stearic acid & tripalmitin/ trimyristin/ trilaurin 154-243nm -60 to -45mV AE 88-95% DL 4.87-10.70% - (72) SA-PEG2k-CS K/ SA-PEG2k-IRQ , tripalmitin,SP C 244-410nm, -20 to -29mV AE 51-59% DL1.87-2.20% Release induced by drug diffusion and SLN skeleton dissolution (80) Tripalmitin or tripalmitin with C8/C10 TG 200-537nm -50 to +29mV AE 31 to >90% Initial burst release ~ 20% followed by sustained release in pH 4 buffer (71, 81-83) Gonadorelin GMO 422nm -21mV AE 50 - 69% Biphasic release, 24.4% release in 6h, sustained release in 12 days (first 2h in 0.1N HCl, followed by pH6.8 buffer) (84) Thymopentin or insulin GMO/GP/GT/G B/stearic acid/ octadecyl alcohol/cetyl palmitate 214-449nm INS AE 33-76% Thy AE 47-76% 4h INS release 33-76% (first 2h in pH 1.2, later in pH 6.8 media); low burst with GP; hydrophobicity & viscosity influence AE and release (85) GP 305nm,-18mV INS AE 57% Thy AE 62% 10% (Thy) & 15% (INS) release at 4h (first 2h in pH 1.2, later in pH 6.8 media) (86) Leuprolide Stearic acid, Sodium Stearate ~ 400nm -46mV AE 28 - 75% DL 0.28-0.76% 8.3% or 23.7% release in 2h in PBS, depending on preparation method (74) Levothyroxine Tripalmitin, palmitic acid 188nm -23mV AE > 99% DL 33.17% Slowed release following initial burst in pH 7.4 medium (75) Lysozyme TG14/TG18/DG /MG - - In fasted state simulated intestinal fluid, lipase-mediated release for TG14, TG18 and DG, not for MG (87) C8/C10 TG: Miglyol 812N (Caprylic/capric triglycerides); CSK: goblet cell ligand CSKSSDYQC; DL: drug loading; DG: diglycerides; GB: glyceryl behenate; GMO: glyceryl monostearate; GP: glyceryl palmitostearate; GT: glyceryl tripalmitate; IRQ: cell-penetrating peptide IRQRRRR; MG: monoglycerides; OAA: octadecyl alcohol; SA-PEG2k: PEG-40 stearate; sCT: salmon Calcitonin; SPC: soybean phosphatidylcholine; TG14: trimyristin; TG18: tristearin. Table 3. Examples of compostions of SLN for oral delivery of peptide/protein drugs: the lipids involved, physicochemical properties, drug entrapment, and in vitro release profile. Chapter 1 45 With regard to the release properties of SLN, the controlled release capacity is highly dependent on composition. For example, Christophersen and co-workers studied the impact of the glycerides chain on the release profile of the model protein, lysozyme (87). It was concluded that a lipase-mediated degradation mechanism was the main trigger for peptide release from SLN composed of triglycerides. Interestingly the peptide delivery profile clearly matched the release profile of free fatty acids. In the case of diglyceride-based SLN, the drug release mechanism involved both a lipolysis mediated and peptide diffusion mechanisms. Finally, the release from monoglyceride-based SLN was found to occur by simple diffusion through the lipid channels created in the imperfect matrix. The study concluded that peptide release rate followed the trend: monoglyceride > diglyceride > triglyceride, suggesting that it is possible to control the protein release by using the adequate mixture of lipids. On the other hand, in other studies, it has been reported that the incorporation of surfactants, liquid-crystal forming ingredients, i.e. glyceryl monostearate (GMO) and the co-entrapment of hydrogels may significantly contribute to the controlled release of the associated protein (66, 67, 84, 86). In addition, to the above-indicated diffusion and degradation mediated mechanism of release, the specific ionic/hydrophobic interaction of the peptide molecules with the lipid components is expected to influence the release profile. As indicated, the cationization of the peptide drug has been a strategy to increase its entrapment due to the ionic interaction with anionic lipids and, such interaction, was found to impact the subsequent release. Interestingly, depending on the nature of the peptide and its ability to interact with the counter-ion, different effects have been observed. For instance, the incorporation of anionic lipid lecithin into SLN led to a gradual release of the cationic peptide sCT (up to 40% in 6h) (71). In a separate study, the incorporation of increasing amounts of lecithin contributed to the sustained release of vascular endothelial growth factor (VEGF) for over 45 days (88). A very different effect was observed for insulin-SC reverse-micelle loaded SLN. In this case, insulin was initially positively charged in an acidic medium in order to promote the entrapment within the SC micelles. However, once the system reached the physiological pH, the charge of insulin was inverted and the repulsion between the drug and bile salts broke down the systems, thus giving rise to burst drug release (73). Chapter 1 46 Figure 3. Peptide release mechanisms based on (A) ionic disassociation followed by diffusion through channels of the lipid matrix, (B) simple diffusion through channels of the lipid matrix and (C) lipase-mediated degradation of the lipid matrix. In conclusion, the SLN formulation approaches used so far have involved the use of counter ions to increase the hydrophobicity of the peptide prior to encapsulation or in the process of encapsulation, the formation of peptide loaded reverse-micelles, as well as the use of the double emulsion approach to facilitate the physical entrapment of the peptide molecules. On the other hand, peptide controlled release can be achieved using lipidic mixtures with lipids owning distinct degradation profiles and also through the control of the porosity of the lipid matrices. 3.1.2. Microemulsions and nanoemulsions Microemulsions are a thermodynamically stable and isotropically transparent dispersion of two or more immiscible liquids stabilized by one or more suitable surfactants (89, 90). Both water-in-oil (w/o) and oil-in-water (o/w) microemulsions may form, although the interpretation is complex, since microemulsions are not true emulsions but colloidal solutions with solubilised water or solubilised oil microemulsions are characterized by an ultra-low A B C Chapter 1 47 interfacial tension between oil and water phases brought about by the selected surfactants and while they may have complex and diverse microstructures varying from micro-droplet, to bicontinuous and solution types, they are generally considered to exhibit a size below 100 nm. Microemulsions can form spontaneously, given the low/limited energy required to produce a thermodynamically stable system. Nanoemulsions, on the other hand, are not thermodynamically stable and require high energy input to produce a kinetically stable system of droplets less than 200nm (91, 92). With regard to the preparation techniques, microemulsions are prepared by mixing the oily components plus the surfactant/s and co-surfactant/s in adequate proportions (93), whereas nanoemulsions may be obtained using high-energy processes (homogenizer, microfluidizers or ultrasonicator) to manufacture emulsions (94-96), or lower-energy approaches including spontaneous emulsification (solvent displacement) or phase inversion temperature (PIT) methods (92, 97-100). The former method successfully avoided high temperature that may hamper the bioactivity of peptides, and the second approach omitted the usage of organic solvent. From drug delivery perspective, micro/nanoemulsions are attractive due to the existence of microdomains of different polarity within the same single phase solution, which can facilitate solubilisation of either hydrophilic or lipophilic materials. The encapsulation of the hydrophilic peptides into micro/nanoemulsions has been achieved through the formation of simple w/o emulsion or multiple w/o/w emulsions (92). Additionally, since medium chain lipids can be more easily mixed with aqueous phase compared to long chain lipids, they are also commonly used in emulsions to facilitate the solubilisation and encapsulation of hydrophilic drugs (101, 102). Finally, the potential interaction of the peptides with specific components, either lipids or surfactants, is expected to significantly contribute to the loading of peptides and their release mechanism. Overall, the efforts dedicated to these peptide formulations, have culminated in over 80% entrapment efficiency for various peptidic drugs such as insulin, sCT and BSA (Table 4). Chapter 1 48 Peptide Lipids and emulsifiers Emulsion Type Size and Z-potential Drug entrapment Ref. Insulin Triacetin; DMAB, propylene glycol W/O 161nm AE 85% (103) Oleic acid; polyglyceryl-6-dioleate, PEG-8 C8/C10 glycerides W/O 108nm, Insulin-CS complex +30mV Insulin complex AE 79% (104) Soybean oil; HCO-60, L-1695, sodium cholate glycerin S/O/W 1.1μm - (105, 106) Soybean oil or Medium chain TG; Cetyl PEG/PPG10/1 dimethicone (W/O); Tween 80 (W/O/W) W/O/W 1-20μm AE> 95% (107) Triolein., oleic acid, DHA, EPA, stearic acid, linoleic acid, linolenic acid, EPC, PEA; Span® 80 (W/O); Tween® 80 (W/O/W) W/O/W - - (108) Soybean oil, triolein, trilinolein, oleic acid, linoleic acid, EPC,PEA; Span® 80 (W/O); Tween® 80 (W/O/W), W/O/W - - (109) Oleic acid, EPC, PEA; Span® 80 (W/O), Tween® 80 (W/O/W), W/O/W 263 to 591nm - (110) Lauric acid, palmitic acid, stearic acid, linoleic acid, palmitoleic acid; SGC W/O - - (111) Insulin & aprotinin Olive oil, C8/C10 TG, PEG-HCO; SPC W/O <200nm - (112) sCT & Aprotinin Liquid paraffin; cholesterol, Arlacel 1689, Atlas SCS 2054 (W/O); Synperonic PE/F127 (W/O/W) W/O/W 14μm AE 92.1% (113) TAT (TAMRA labelled) C8/C10 TG; Polyethoxylated castor oil W/O 21.4nm - (8) BSA Isopropyl myristate; polyethoxylated castor oil, propylene glycol, Tween®20 W/O 21.8nm, -24.8mV AE > 90% (114) EFE-d C8/C10 TG; Caprylocaproyl macrogol-8 glycerides; Polyglyceryl-3 dioleate W/O 6.86nm - (115) Chapter 1 49 SK&F106760 Captex355; Capmul® MCM, Tween® 80 W/O 15.2nm - (116) Arlacel 1689: sorbitan oleate (and) polyglyceryl-3 polyricinoleate; Captex355: glyceryl tricaprylate/tricaprate; Capmul® MCM: mono-diglyceride of medium chain fatty acids (mainly caprylic and capric); DHA: docosahexaenoic acid; DMAB: didoceyldimethylammonium bromide; EFE-d: earthworm fibrinolytic enzyme; EPA: eicosapentaenoic acid; EPC: egg phosphatidylcholine; HCO-60: polyoxyethylene hydrogenated castor oil; L-1695: sucrose lauric ester; PEG-HCO: PEG-7 hydrogenated castor oil; PEA: phosphatidylethanolamine. Table 4. List of emulsion, microemulsions and nanoemulsions for oral delivery of p/p drugs: lipids involved, emulsion type, physicochemical properties and drug entrapment. There are very few studies on the mechanism of release of peptides from micro/nanoemulsions. Overall, as in the case of the loading, the specific interaction of the peptides with lipid constituents and surfactants is expected to influence the partition between the oily phase and the external aqueous medium. In addition, the degradation of the oil in external media is expected to trigger drug release. In a study with emulsions that contains medium chain triglycerides/soybean oil, surfactants Tween 80® and cetyl PEG/PPG-10/1 dimethicone, the authors found an osmotically-driven swelling effect that caused the breakdown of the emulsions and further leads to drug release (107) (Figure 4). In other studies, the authors proposed to avoid undesired peptide release by protecting the emulsions from degradation (112, 113). Moreover, from a product development perspective, there is the possibility to incorporate emulsions, either in a liquid or dry form in pH-sensitive pharmaceutical dosage-forms (105, 106). All these activities on micro/nanoemulsion-based formulations has led to significant advancements for enhancing oral peptide delivery (30, 117-119). Further promising development in microemulsions also has given rise to self-emulsifying drug delivery systems (SEDDS, SMEDDS and SNEDDS) and similar structures, as described below. Chapter 1 56 BSA Glycerol trioleate PBCA 160nm DL(%): 1-4 Higher pH / loading and lower drug MW promote burst release; surface erosion induced drug release; lower MW protein diffuse through PBCA wall easily (148) D-Lys6-GnRH Ethyl oleate PECA 190nm, +8.8mV 98.6 <5% release at 6h in GIT media; 60% release in plasma (161) Carbopol: Carbopol®940, cross-linked polyacrylate polymer; C8/C10 TG: caprylic/capric triglycerides; DL: drug loading, PBCA: poly(butyl-cyanoacrylate); PECA: poly(ethyl-cyanoacrylate); TDI: tolylene 2,4-diisocyanate. Chapter 1 57 Table 5. Nanocapsules-based oral peptide delivery formulation indicating lipids cores and polymer shells composition, physicochemical properties, drug entrapment and in vitro release profile. 3.1.5. Liposomes Liposomes are vesicles with a diameter from 10nm to more than micron, comprising a well-defined aqueous core and one or more amphiphilic bilayers made of phospholipids and cholesterol. In addition, some polymer compositions incorporate in their structure different surfactants and polymers (162, 163). These vesicles are by far the most extensively studied vesicular system since its discovery in 1961 (164, 165) and, so far, this research activity has been translated into 13 FDA approved liposomes-based products for human use, containing low molecular weight drugs and intended for injection. This prior knowledge has been fundamental for the development of liposomal formulations intended for oral peptide delivery. Other liposome-like nanostructures, such as niosomes and archeosomes, have also been proposed for the oral administration of peptides and proteins (166). In most cases, peptide-loaded liposomes have been prepared by the film hydration (167-183), and reverse-phase evaporation methods (184-192), followed by particle size reduction treatments including sonication, extrusion or high pressure homogenization. To date numerous peptides and proteins such as insulin, sCT, albumin, adamantyltripeptides, globulin, leuprolide and others have been entrapped into liposomes, with the final goal of enhancing oral bioavailability (193). Table 6 shows representative examples of compositions reported in the last years for oral peptide delivery. Peptides can be entrapped into the aqueous core of the liposomes. However, as described for other lipid-based nanostructures, the association efficiency of peptides within liposomes is also greatly influenced by its ionic/hydrophobic interaction with the liposomal components and also by the rigidity of the by-layer formed (179, 180). Chapter 1 58 Up to now, over 90% association efficiency of insulin and sCT has been successfully achieved, making use of a variety of lipids including, diplamitoyl-phosphatidylcholine (DPPC), distearoyl-phosphatidylcholine (DSPC), stearylamine, dicetyl phosphate, or blends of these lipids with cholesterol. For example, when DPPC was mixed with higher phase transition temperature lipids like dipalmitoyl-glycerophosphoethanolamine (DPPE) to form liposome, the AE of insulin was significantly enhanced (175). It has also been shown that an increase in the concentration of lipids in liposomes commonly leads to a higher entrapment of peptides, while the absolute drug payload is inevitably decreased simultaneously. Finally, the association other biomaterials, i.e. lectins, to the surface of the liposomes, has sometimes resulted in an increase in the entrapment efficiency (167). As described for other nanocarriers, the peptide loading capacity of peptides into liposomes is limited by the capacity of the aqueous core and the lipid bi-layer to hold the peptide molecules. In some cases, it has been found that an increase in the amount of peptide to be loaded may lead to a destabilization of the nanostructure (173, 175, 182). Therefore, it is important to carefully design the composition and structure of the liposomes in order to reach the maximum loading, while maintaining the stability and controlled release properties of the system. The peptide release rate from liposomes is known to be highly influenced by liposomal fluidity, which depends on the phase transition temperature of phospholipids involved. In particular the use of long and hydrophobic phospholipid chains, as well as the presence of cholesterol are known to enhance the rigidity of the liposomes wall and, hence, their colloidal stability and capacity to control the release (194). Additionally, the use of special lipids, such as diether or tetraether lipids – the resulting liposomes were named as archeosomeswere reported to maintain the structural integrity against extreme pH, to prevent the degradation by bile salts and lipases and to help the control of the release (170). Similarly, the use of Chapter 1 59 polyoxyethylene alkyl ethers - the resulting nanostructures were named as niosomeshave led to a significantly delayed insulin release (171). More drastic procedures for enhancing the stability and controlling the release of peptides from liposomes have relied on their association with other biomaterials and nanostructures. For instance, a formulation of thiolated chitosan-coated niosomes exhibited a 12% insulin release in SGF in 5h (189), a result that was attributed to the protection of the vesicle structure in the GI environment. On the other hand, liposomes coated by silica nanoparticles exhibited an improved stability and a sustained release profile for the insulin (169). Our group also reported that the complexation of liposomes with insulin-loaded CS nanoparticle resulted in an enhancement of the colloidal stability of CS nanoparticles with the possibility to tailor the insulin release profile (174, 195). Finally, the encapsulation of the insulin-loaded liposomes into larger liposomes (196), or the surface PEGylation of liposomes were adequate strategies towards this goal (214). Overall, efficient peptide incorporation in the inner aqueous core of liposomes has been achieved, although the loading capacity of these nanocarriers clearly depends on the interaction of the peptide molecules with the lipids, surfactants and potential polymers involved in the process. In general, it is accepted that a substantial amount of peptide is associated to the by-layer, rather than encapsulated into the core, and this often leads to the unpredicted leakage of the peptide. However, an increasing number of approaches are being disclosed with the final goal of a tailor-made design of liposomes for oral peptide delivery. Currently, some liposomal drugs for oral peptides delivery are in clinical trials, like clinical phase II drug HDV-I (Diasome pharmaceuticals, USA) for oral insulin delivery. Chapter 1 60 Peptide System Size & Z-pot AE (%) In vitro release profile Ref. Insulin Biotin-modified liposomes (SPC) 85-309nm 22-65 - (184, 185) Bile salts added Liposome (SPC) 150nm (SGC) 30 In 6h, 57.7% release in SGF; 63.8% release in FaSSGF (186-188) Trimethyl CS coated niosome (Span®60) 100-180nm (non-coated) 75 12% release in SIF by 5h (189) Folic acid added multi-coat liposome (PC, SAM) 266nm +25mV 93 < 20% release in SGF, SIF & pH 7.4 PBS in 2h; up to 72% release in PBS in 24h (168) Silica nanoparticle coated liposome (DPGPC) 255-297nm ~-15mV ~ 70 Negligible release 45min in SGF; slowly and continuous release for up to 8 h; slower release for coated liposomes (169) Archaeosome (tetraether lipids) 210nm -36mV 18.6 70% and < 50% release in 4h in artificial intestinal and gastric condition, respectively (170) Niosome (BrijTM,DCP) 4.2-13.4μm 8.7-41.5 26.3% insulin released during 24h in SIF (171) Fusogenic liposome (EPC, PG) - 3.3 - (197) Lectin modified liposome (SPC) 190-194nm +3.4 to +8.7mV 40-83 - (190) PEG-2000 or mucin coated liposome (DPPC, DSPE-PEG) 453-479nm 35-38 Surface coating resulted in resistance to bile salt digestion and slow release in GI tract (172) Aprotinin added double liposomes (SPC, SA/PS) 8-9μm 7.6-23 Double liposome show slightly slower release (196) Complex of liposome & CS nanoparticle (DSPC, DPPS) 1.5-3μm -17.7mV - Lipid coat induced delayed drug release (174) Liposome (PEtOH) 50-250nm > 80 - (175) WGA-carbopol modified liposome (DSPC, SAM) 187-201nm -44 to -57mV 91 - (176) sCT Thiolated CS coated liposome (DPGPC, DPPE-MCC) 605-709nm +28 to +44mV 69 No evident release in SGF; sustained release in SIF with less than 20% burst (198, 199) Pectin-liposome nanoxomplexes (DSPC, SAM) Below 1μm -40 to -60mV 49.7 - (177) Chapter 1 61 CS coated liposomes (DSPC, DCP) 305nm-1.8μm 100 (20% EE) - (178) CS coated liposome (DSPC, DCP) with different size 473nm-4.1μm +30 to +35mV > 90 Sustained release at pH 6.8 (179, 180) Double liposomes (DMPC, DPPC, SAM, DPPG) 2.1μm 36.2-62.9 - (196, 200) Albumin CS coated liposome (SPC or EPC) 128nm +5.4mV 50% Drug : lipid ratio 2.5% - (182) Octreotide Archeosome (tetraether lipids) 130-207 nm Drug : lipid ratio 1.2 to 13% - (183) Leuprolide PSCG coated liposome (PC) 112-168nm -8.8 to -20mV 29.4-37.1 Drug release due to vesicle disruption (167) Epidermal growth factor Liposome (DMPG, DOPC, triolein) 1.4μm Up to 60 Sustained release; 47% release in SGF and 35% release in SIF in 6h (191) BrijTM: polyoxyethylene alkyl ethers; DCP: dicetyl phosphate; DMPC: dimyristoyl-phosphatidylcholine; DMPG: dimyristoyl-phophatidyl-glycerol; DOPC: dioleoyl-phosphatidyl-choline; DPPC: dipalmitoyl-phosphatidyl-choline; DPPCG: dipalmitoyl-glycero-phosphocholine; DPPE-MCC: dipalmitoyl-glycero-phosphoethanolamine-maleimidomethyl-cyclohexane-carboxamide; DPPG: dipalmitoyl-phosphatidyl-glycerol; DPPS: dipalmitoyl-phosphatidyl-serine; DSPC: distearoly-phosphatidyl-choline; DSPE-PEG: distearoylphosphatidylethanolamine-poly-(ethylene glycol) 2000; FaSSGF: fasted state simulating gastric fluid, PC: phosphatidylcholine; PEtOH: phosphatidylethanol; PG: L-dimyristoyl phosphatidylglycerol; PS: phosphatidylserine; PSCG: O-palmitoylscleroglucan; SAM: stearylamine. Table 6. List of vesicular systems for oral delivery of p/p drugs: lipids involved, physicochemical properties, drug encapsulation and release profile. Chapter 1 62 3.2. Interaction of lipid-based nanocarriers with the biological environment Given that lipids are biodegradable, and in particular susceptible to intestinal lipase mediated degradation, a comprehensive understanding of the behaviour of the lipidic nanocarriers upon contact with the gastro-intestinal milieu is needed in order to reliably predict in vivo performance. This degradation of the lipid materials may also be followed by the degradation of the associated peptides. Finally, the transport of these nanocarriers across the mucus barrier and subsequent interaction/transport across the intestinal epithelium are expected to have important consequences on the overall performance of the nanocarriers. In the following sections, we will report the knowledge accumulated so far with regard to the ability of lipid-based nanocarriers to deal with these biological barriers. 3.2.1. Stability in simulated biological fluids As indicated in section 2.1 (table 1), selected lipid excipients, namely long chain glycerides and fatty acids, may confer lipid-based nanocarriers with a resistance against enzymatic degradation (5, 6). In the case of SLN, the solid physical state of the lipids is expected to help protect the encapsulated peptide (48, 77, 78, 85, 87). Similarly for NC, the polymeric shell may offer protection from exposure to low gastric pH and enzymatic degradation (81, 83, 130, 135, 145, 152-157). PEG surfactants have also been employed as a general strategy to protect microemulsion systems from aggregation and degradation (8, 104, 107). Finally, the colloidal stability of liposomes and their capacity to protect the encapsulated drug has been achieved through adequate lipid selection (170, 183, 197), the inclusion of enzyme inhibitors (186-188, 196), or the attachment of a protective polymer around the phospholipids bilayer (168, 178). In terms of selecting lipid excipients to protect from enzymatic degradation, it appears that long chain fatty acids are favoured. For example, in a study it was shown that the stability of SLN and their subsequent capacity to protect the encapsulated Chapter 1 63 molecules followed the ranking: tripalmitin (C16) > trimyristin (C14) > trilaurin (C12) (72). In a separate study involving liposomes, it was found that the use of phospholipids with a high phase transition temperature leads to the formation of a gel membrane, which helps in preventing their colloidal stability as well as the stability of the encapsulated peptide Apart from the mentioned triglycerides and phospholipids, there are other lipid materials, which may significantly contribute to the colloidal and molecular stability of lipid-based nanocarriers. Bile salts, in addition to their penetration promoting capacity, have been reported as efficient stabilizers. As an example, insulin loaded-liposomes containing sodium glycocholate (SGC), sodium taurocholate (STC), or sodium deoxycholate (SDC) were reported to have an improved stability in simulated gastrointestinal fluids (186-188). SGC-containing liposomes displayed the highest drug protection and longest residence time. More importantly, this improved stability profile was translated into an enhanced in vivo efficacy in terms of more pronounced glucose level decrease and sustained drug action. As indicated above, the presence of a polymer coating around the liposomes, SLN or NC, can positively influence the stability of these nanocarriers. A common approach involved the use of PEG derivatives and, in particular PEG stearate, which has been used to improve the stability of SLN and the entrapped peptides in gastric-intestinal fluids (75, 80). Similarly, the PEGylation of liposomes with PEG-2000 was reported to enhance the resistance to digestion (172). An alternative approach has made use of the polysaccharide chitosan and PEG-Chitosan in combination with lipids. Different type of nanocarriers, including liposomes and NC, combined with chitosan were found to exhibit an improved stability and overall in vivo performance (81, 135, 139, 174, 195). Chitosan coated tripalmitin SLN maintained particle size upon contact with simulated gastric-intestinal media (71). Chapter 1 64 A drastically different strategy oriented to increase the stability of the encapsulated peptide has consisted in involved the formation of conjugates of peptide/protein with lipids. In particular, fatty acids have been linked to peptidic drugs (sCT, insulin, enkephalin, INF-α, octreotide, tetragastin, demopressin, etc.), the result of which was an improvement of the peptide stability against enzymatic degradation (201-204). Although not systematically investigated, the increased in the fatty acids carbon number (from 2 to 5) in fatty acids was found to positively impact the stability of the resulting peptide-lipid conjugate (205). This could be, in some cases, related to the capacity of the amphiphilic conjugates to form micelles (206, 207). Palmitic acid (C16) is the most studied molecule for this peptide lipidization approach. Overall, it can be concluded that the appropriate selection of lipids with different properties and/or the addition of auxiliary excipients, such as surfactants and polymers, to lipid-based nanocarriers has been successfully employed to overcome the harsh conditions present in the GI environment. However, given that these studies have been generally performed using simulated gastric-intestinal fluids in the presence of enzymes, the extrapolation of this in vitro data to the in vivo conditions remains unclear. 3.2.2. Interaction with the mucus layer In theory, due to their innate hydrophobicity, lipid-based nanocarriers are not supposed to optimally diffuse across the water-rich mucus layer. However, the combination of hydrophobic lipids with amphiphilic lipids, surfactants and polymers offers the possibility to modify the muco-diffusion of lipid-based nanocarriers. Traditionally, the use of bioadhesive polymers and, in particular chitosan and its thiolated version (71, 139, 208), alginates (141), pectin (177), and acrylic polymers (143, 176, 180), have been proposed as a strategy to favour the interaction of the nanocarriers with the mucus, with a view to prolonging residence time in the intestinal Chapter 1 65 tract. However, more recently, it has been recognised that increasing the muco-interaction may not be an optimal strategy for nanocarriers to penetrate to the underlying epithelium. In fact, the nanocarriers may be retained in the mucus layer and be eliminated via mucus turnover. On the contrary, the emphasis is now oriented towards achieving an adequate mucodiffusion. It is generally considered that a small nanocarrier size, a neutral or negative surface charge, and enhanced surface hydrophilicity are favourable properties for achieving adequate muco-permeation (209, 210). This postulate has been confirmed for a number of lipid-based nanocarriers, including SLN (80), SNEDDS (211, 212), NC (129) and liposomes (179, 186-188). Clearly, the surface characteristics of the nanocarrier can be tailored to optimise muco-diffusion. The potential of protecting the nanocarriers with a hydrophilic layer made of PEG, is particularly noteworthy given the ability to enhance diffusion through the hydrophilic mucous layer and also prevents interaction with biological enzymes and mucin. The first evidence of this positive role of PEG was reported by our group in the late 90’s (213) and has been extensively validated thanks to the work by Hanes and co-workers (210, 214-216). Mechanistic studies demonstrating the benefit of PEG in the mucodiffusion of lipid-based nanocarriers has been shown for NC containing PEG stearate (129), SNEDDS involving PEG oleate (211) and liposomes modified by poloxamer 407 (217). Finally, it is important to keep in mind that the concentration of nanocarriers interacting with the mucus layer may significantly alter their mucodiffusion behaviour (216, 218). While our understanding of nanocarrier interaction with the mucous layer has been greatly improved as a result of these studies, there remains a need for a more systematic study of the properties influencing muco-adhesion, muco-diffusion and stability of nanocarriers in the intestinal environment. Additional efforts should focus on identifying experimental techniques illustrating the intestinal mucus environment Chapter 1 72 and is also believed to be sensitive to absorption enhancers (12, 233). The merits of lipid excipients, in terms of enhancing permeability via the colonic mucosa have been well established (234, 235). Table 7 presents a summary of studies involving colon targeted local delivery of oral peptides, where lipid excipients are involved as a component in these multifunctional drug delivery systems, as mentioned in section 2.1. However, despite significant efforts to develop site-specific targeting of peptides, most notably insulin, developing drug delivery system with demonstrate reliable and consistent systemic absorption from the large intestine remains problematic, as recently reviewed (236, 237). In summary, though the technologies described here suggest the potential of lipid-based nanocarriers for oral peptide delivery, further mechanistic and bio-relevant studies are required in order to get a comprehensive knowledge on the complex interactions that occur in vivo between peptides, lipid excipients and the intestinal milieu. This mechanistic knowledge will help understanding the in vivo performance of the nanocarriers. Peptide Lipid excipient (PE) Delivery system Additional functional materials PK/PD observations Ref. Insulin SGC Enteric coated capsules HPMC phthalate (coat), chitosan (capsule shell) 5.73% BA (rats) Inclusion of SGC increased PA to 3.5% (238) Sodium oleate Enteric coated microspheres Bacitracin (PI), hydroxypropylethyl cellulose acetate succinate (coat) Hypoglycaemic effect observed between 2-7h post dosing (rats) (239) SGC Coated tablets (CODES®) Lactulose (core), Eudragit E, HPMC, Eudragit (coat), Camostat mesilate (PI), EDTA (PE) Absolute BA of 0.34% by inclusion of SGC from 0.13% (not significant). Significantly reduced glucose levels between 6-8h post dosing (dogs) (240) Chapter 1 73 Oleic acid decyl ester, PEG-8 C8l/C10 glycerides, polyglycerol oleate Enteric and time-controll ed release coated capsules Eudragit S, NE30D and cellulose acetate phthalate (coat), Carbosil (gelling agent) aprotinin (PI), silicium dioxide (gelling agent) PA increased to 6.2% from 2.1% (dogs) (50) Calcitoinin SGC Enteric coated capsules HPMC phthalate (Coat) chitosan (capsule shell), bacitracin, aprotinin (PI), S-Nitro-N-acetyl-penicillamin e (PE) PA increased to 6.3% from 0.04% for control (rats) (241) Sodium Caprate Minispheres (Smpill®) Caprylic/capric/linoleic TG, Cremophor EL (microemulsion), gelatin (core) Absolute BA of 22.3% after intracolonic administration from 7.0% (rats) (242) Ciclosporin (CsA) Caprylic/Capric Triglyceride, Cremophor EL Enteric and time controlled release coated minispheres (Smpill®) Ethyl cellulose/pectin (coat), gelatin (core) Increased colonic delivery of CsA Increased CsA levels in colonic tissue (243) Erythropietin (EPO) PEG-8 capryl/caprylic acid glycerides Drug/lipids adsorbed on solid adsorbent (carbon nanotubes) Casein, lactoferrin, Explotab Absolute BA of 11.5% after intra-jejunal administration (rats). >6 fold BA compare to non-lipid control (244) Heparin (LMWH) Capric acid (C10) Enteric coated capsules (GIPET I®) Eudragit (coat) Absolute BA of 3.9-7.6% (humans) (245) Desmopressin Microemulsions Enteric coated capsules (GIPET II®) Eudragit (coat) monoand di-glycerides of caprylic an capric acid Absolute BA of 2.4% from 0.2% for control in humans (245) Octreotide Sodium caprylate Enteric coated capsules (Octreolin®) Oily suspension including polyvinyl pyrrolidone, polysorbate 80, glyceryl monocaprylate, glyceryl tricaprylate, and magnesium chloride Relative oral BA of was similar to SC injection in humans. Achieved efficacy in controlling IGF-1 and GH up to 13 months (13, 14) PI: protease inhibitor; PE: permeation enhancer. Chapter 1 74 Table 7. Colonic local peptide/protein delivery using lipid excipients 5. Conclusion and future perspectives In summary, an appropriate design and development of optimal lipid-based nanocarriers requires consideration of the distinct merits of various formulations and the peptide or protein characteristics. Using a range of lipid based nanocarriers systems, adequate peptide entrapment has been accomplished and certain formulations have shown ability to control the drug release. It is also increasingly recognised, that in order to overcome physiological, biochemical and biopharmaceutical barriers to delivery, a multifunctional drug delivery system employing protease inhibitors to enhance stability, permeation enhancers to enhance uptake and modified release technology for site specific delivery. Critically, the roles of lipid excipients will therefore hold significant promise. To date, good entrapment, adequate protection (especially against enzymatic degradation), prolonged retention in GIT and enhanced permeation have been demonstrated employing lipid based nanocarriers. However, more mechanistic studies are needed to shed new insights on the understanding of the interplay between nanocarriers, peptides and physiological conditions in the intestine. It will also be helpful to create databases to summarise the existing data as well as to have uniformed model for both in vitro and in vivo tests. Meanwhile, the in vitro - in vivo bio-relevance must receive more attention and the adequate correlation could be crucial for the robust development of commercially viable products. From the formulation point of view, the combination of drug and vehicle oriented approaches, i.e. lipidization of peptide prior to encapsulation into the nanocarriers, maybe crucial. 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Promoting absorption of drugs in humans Chapter 1 96 using medium-chain fatty acid-based solid dosage forms: GIPET. Expert Opin Drug Deliv. 2006;3(5):685-92. Antecedentes, Hipótesis y Objetivos Background, Hypothesis and Objectives 97 Antecedentes, Hipótesis y Objetivos Background, Hypothesis and Objectives Antecedentes, Hipótesis y Objetivos Background, Hypothesis and Objectives 104 Objectives Bearing in mind the background information and hypothesis described above, the objective of this thesis has been to rationally design and develop two different nanocarriers for the oral administration of peptides. The systematic characterization of these two different nanocarriers and the evaluation of their mechanistic behavior both, in vitro and in vivo, has been expected to contribute to the understanding of the interaction of nanocarriers with the biological barriers associated to the oral modality of administration. More precisely, from an experimental viewpoint, the following activities have been undertaken: Polyarginine nanocapsules: 1. Development of new compositions of PARG with a nanometric size, a narrow size distribution, and an adequate stability in simulated biological fluids. 2. Determination of the capacity of the nanocapsules to load and control insulin release. For this, a study of the key factors that influence the entrapment of the peptide has been performed. 3. Study of the in vitro toxicity and mechanism of interaction of the nanocapsules with the intestinal epithelium (Caco-2 model cell line and human intestinal epithelium). For this, the capacity of the nanocapsules to enhance the permeability and the transport of insulin has been evaluated. This capacity was compared to that of the free PARG. Antecedentes, Hipótesis y Objetivos Background, Hypothesis and Objectives 105 4. In vivo studies aimed at understanding the interaction of fluorescence-labelled PARG NCs with the intestinal epithelium after oral administration, and also at determining the bioactivity of insulin-loaded nanocapsules following subcutaneous and intra-intestinal injection to healthy rats. Nanocomplexes: 1. Development of complexes of insulin and polymer A and polymer B with a nanometric size, a narrow size distribution, and high insulin association efficiency. 2. Study of the stability of the r colloidal stability of the nanocomplexes in simulated intestinal fluids and under storage condition, enhance their capacity to protect the entrapped peptide drug from enzymatic degradation, and to optimize their muco-interaction profile. 3. Study of the in vitro toxicity and mechanism of interaction of the nanocomplexes with the intestinal epithelium (Caco-2 model cell line and human intestinal epithelium). 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Chitosan Nanocapsules as Carriers for Oral Peptide Delivery: Effect of Chitosan Molecular Weight and Type of Salt on the In Vitro Behaviour and In Vivo Effectiveness. Journal of Nanoscience and Nanotechnology. 2006;6(9):2921-8. 7. Prego C, Torres D, Fernandez-Megia E, Novoa-Carballal R, Quinoa E, Alonso MJ. Chitosan-PEG nanocapsules as new carriers for oral peptide delivery. Effect of chitosan pegylation degree. J Control Release. 2006;111(3):299-308. 8. Tobio M, Sanchez A, A. Vila IS, Evora C, Vila-Jato JL, Alonso MJ. The role of PEG on the stability in digestive fluids and in vivo fate of PEG-PLA nanoparticles following oral administration. Colloids and Surfaces B: Biointerfaces. 2000;18:315-23. 9. Tobío M, Gref R, Sánchez A, Langer R, Alonso MJ. Stealth PLA-PEG Nanoparticles as Protein Carriers for Nasal Administration. Pharmaceutical Research.15(2):270-5. 10. Yamaki T, Kamiya Y, Ohtake K, Uchida M, Seki T, Ueda H, et al. A mechanism enhancing macromolecule transport through paracellular spaces induced by Poly-L-Arginine: Poly-L-Arginine induces the internalization of tight junction proteins via clathrin-mediated endocytosis. Pharm Res. 2014;31(9):2287-96. Antecedentes, Hipótesis y Objetivos Background, Hypothesis and Objectives 107 11. Ohtake K, Maeno T, Ueda H, Ogihara M, Natsume H, Morimoto Y. Poly-L-arginine enhances paracellular permeability via serine/threonine phosphorylation of ZO-1 and tyrosine dephosphorylation of occludin in rabbit nasal epithelium. Pharm Res. 2003;20(11):1838-45. 12. Miyamoto M, Natsume H, Satoh I, Ohtake K, Yamaguchi M, Kobayashi D, et al. Effect of poly-l-arginine on the nasal absorption of FITC-dextran of different molecular weights and recombinant human granulocyte colony-stimulating factor (rhG-CSF) in rats. International Journal of Pharmaceutics. 2001;226(1–2):127-38. 13. Ohtake K, Natsume H, Ueda H, Morimoto Y. Analysis of transient and reversible effects of poly-l-arginine on the in vivo nasal absorption of FITC-dextran in rats. Journal of Controlled Release. 2002;82(2–3):263-75. 14. Nemoto E, Takahashi H, Kobayashi D, Ueda H, Morimoto Y. Effects of Poly-L-arginine on the Permeation of Hydrophilic Compounds through Surface Ocular Tissues. Biological and Pharmaceutical Bulletin. 2006;29(1):155-60. 15. Rawat A, Yang T, Hussain A, Ahsan F. Complexation of a Poly-l-Arginine with Low Molecular Weight Heparin Enhances Pulmonary Absorption of the Drug. Pharmaceutical Research. 2007;25(4):936-48. 16. Ong JTH, Stetsko G, Jennings R. Novel methods and compositions for enhanced transmucosal delivery of peptides and proteins. Google Patents; 2005. 17. Nemoto E, Ueda H, Akimoto M, Natsume H, Morimoto Y. Ability of Poly-L-arginine to Enhance Drug Absorption into Aqueous Humor and Vitreous Body after Instillation in Rabbits. Biological and Pharmaceutical Bulletin. 2007;30(9):1768-72. 18. Lozano MV, Lollo G, Alonso-Nocelo M, Brea J, Vidal A, Torres D, et al. Polyarginine nanocapsules: a new platform for intracellular drug delivery. Journal of Nanoparticle Research. 2013;15(3). Chapter 2 108 Chapter 2 109 Chapter 2 Polyarginine Nanocapsules: a Potential Carrier for Oral Peptide Delivery This work is done in collaboration with: Federico Benetti 1, Ana Beloqui 2, Aloïse Mabondzo 3, Ilaria Marigo 4, Patrik Lundquist 5, Sulay Tovar 6, Véronique Préat 2, Per Artursson 5, Carlos Dieguez 6 1. ECSIN-European Center for the Sustainable Impact of Nanotechnology, Veneto Nanotech S.C.p.A., 45100, Rovigo, Italy. 2. Louvain Drug Research Institute, Advanced Drug Delivery and Biomaterials, Université Catholique de Louvain, B-1200, Brussels, Belgium. 3. Biology and Technology Institute of Saclay (iBiTec-S), CEA Life Sciences Division, 91191, Gif sur Yvette Cedex, France. 4. Veneto Institute of Oncology, IOV-IRCCS, 35128, Padova, Italy. 5. Department of Pharmacy, Uppsala University, Box 580, SE-751 23 Uppsala, Sweden 6. Biomedical Research Group, Center for Research in Molecular Medicine and Chronic Diseases, University of Santiago de Compostela, 15782, Santiago de Compostela, Spain. Chapter 2 110 Chapter 2 111 Abstract The aim of this work has been to rationally design and characterize a new type of nanocapsules composed of an oily core and a polyarginine (PARG) shell intended for oral peptide delivery. PARG, a cationic polyaminoacid, was selected based on its known penetration enhancing properties. After the screening of a number of formulation conditions, a composition containing oleic acid and sodium deoxycholate (SDC) as additional penetrating enhancers, was defined and characterized. The NCs, prepared by a mild solvent displacement technique, exhibited an average size of 180 nm, a low polydispersity (0.1) and high insulin association efficacy (AE). Another relevant feature of these NCs was their stability upon incubation in simulated intestinal fluids (SIF, FaSSIF-V2, FeSSIF-V2), in the presence of lipase and pancreatic enzymes. They were also stable during long-term storage (over 45 days). Moreover, studies on Caco-2 cells indicated that the NCs exhibit a concentration-dependent cytotoxicity, whereas no evident toxicity was found on human intestinal tissue. With regard to their mechanism of action, the results showed that PARG NCs led to a reversible reduction of the TEER of Caco-2 monolayers, which might be responsible for the observed facilitated transport of the associated insulin (3.54% as compared to 2.73%, observed for insulin mixed with PARG in solution). Finally, in vivo studies performed in rats showed that insulin bio-activity is not depraved during the NC preparation. Overall, this work shows that PARG NCs fulfill specific requirements that make them an attractive vehicle for oral peptide delivery. Chapter 2 112 Chapter 2 113 1. Introduction Diabetes is a chronic disease that threatens the life of human beings. There are 382 million of diabetes sufferers worldwide today and an alarming number of 471 million is supposed to be reached by 2035 (1). Insulin has been widely used to treat this disease via subcutaneous injection. However, this route of administration is not the best option due to the pain and stress caused by repetitive injections. The oral route of administration is generally considered as the best alternative thanks to the good patient compliance and the physiological pathway followed by the peptide after intestinal absorption. Despite these clear advantages, the delivery of peptide molecules by the oral route represents a huge challenge due to the important biological barriers that the peptide molecules need to overcome. Firstly, the labile peptide should be able to resist the harsh environment of the GIT. Secondly, the peptide should cross the mucus layer and the underlying epithelium (2). As a consequence of these critical barriers the oral bioavailability of current peptides administered orally is below 1%. The approaches investigated so far to improve the absorption of peptides administered orally, include the chemical modification of the peptide molecule, the co-administration of protectors or permeation enhancers and the formulation of adequate nanocarriers (3, 4). Nanocarriers such as polymeric nanoparticles, liposomes, micelles, solid lipid nanoparticles, microemulsions, self-emulsifying drug delivery systems, nanoemulsions and NCs have attracted much attention over the past years (5-14). Within this frame, our laboratory has contributed with the design of NCs consisting of an oily core surrounded a shell made of chitosan (7, 15). These NCs allowed us to encapsulate the peptide salmon calcitonin (sCT) and the resulting nanocomposition was found to facilitate and prolong sCT absorption following oral administration. Chapter 2 120 2.6. In vitro release profile of insulin from PARG NCs The in vitro release profile of insulin from PARG NCs was evaluated in both, SIF and FaSSIF-V2 media in three replicates. Specifically, 5 aliquots of 0.25mL of insulin-loaded NCs from the same batch were diluted with 1.25 mL of the desired intestinal media and placed in an incubator at 37oC under horizontal shaking (300 rpm). After 0, 2, 3, 3.5 and 4 hours the different samples were ultracentrifuged (82,656g; 4ºC; 1h) and the free insulin present in the undernatant was quantified by HPLC. Additionally, the insulin present in the NCs cream was evaluated by HPLC through the degradation of the cream with Triton™ X-100, acetonitrile and 0.1% TFA (same method as section 2.4), in this case, the released insulin was calculated by 2.7. Caco-2 cells culture Caco-2 cells were grown in DMEM high glucose with L-glutamine supplemented with 10% heat inactivated fetal bovine serum, 1% Penicillin (100 U/mL), streptomycin (100 μg/mL), and 1% NEAA solution. Cells were maintained at 37 °C in a humidified incubator supplied with 5% CO2. 2.8. Toxicity studies on Caco-2 cells For cytotoxicity evaluation, Caco-2 cells were seeded in 96-well plates at the density of 10,000 viable cells/well, and incubated 24h to allow cell attachment. Cells were then incubated with increasing concentrations of the tested samples. The cytotoxicity of both blank and insulin loaded PARG NCs was determined by measuring metabolic activity with MTS assay and neutral red uptake (NRU). Lactate dehydrogenase (LDH)-based cytotoxicity assay was also used to measure LDH released into media from damaged cells as a biomarker for cellular cytotoxicity and cytolysis. In the study, the cell culture medium was replaced by the blank or insulin loaded PARG NC suspension in cell culture medium at concentration 0.1855, 0.371, 0.742, 1.484, 2.968, Chapter 2 121 4.452 and 5.936 mg/mL. The plate of Caco-2 monolayers were transferred to a humid incubator at 37 ºC with 5% CO2. After 2h incubation, the tested samples were removed. The cells were rinsed with PBS and incubated at 37 °C for 3h with 120µL of fresh culture medium containing 20% MTS solution. Cellular supernatants were then transferred into a new 96-well plate and the amount of soluble formazan produced by cellular reduction of MTS was determined recording absorbance at 490 nm with Synergy 4 microplate reader (BioTek Instruments, Inc., Winooski, USA). Before performing MTS assay, 50 µL of cell culture media were transferred into a new 96-well plate, mixed with 50 µL of working reagent for LDH detection and incubated for 20 min at room temperature in the dark. Reaction was blocked by adding 25 µL of stop solution, and the amount of produced formazan was measured recording absorbance at 500 nm with Synergy 4 microplate reader. Cytotoxicity of PARG NCs was also evaluated by neutral red uptake (NRU) assay. Treated Caco-2 cells were rinsed with PBS and incubated for 3h at 37 °C with 100 µL of cell culture medium containing 10% Neutral Red solution. After incubation, medium was removed and cells rinsed twice with Dulbecco’s PBS before adding Neutral Red Assay solution. Plate was shaken 45 minutes at room temperature, and absorbance recorded at 540 nm with Synergy 4 microplate reader. The cytotoxicity of a control NE (without PARG shell) and also that of the PARG polymer solution was evaluated by NRU assay. In the study, the NE was tested at same concentration as the PARG NCs, while the PARG polymer solution was tested at concentration 0.009, 0.017, 0.034, 0.068, 0.137, 0.205 and 0.274 mg/mL, which are the amounts of PARG polymer involved in the tested PARG NCs concentration successively. The assays were repeated three times independently, each run as three independent technical replicates. Results are reported as a percentage of control and expressed as mean ± standard deviation. Data from in vitro testing were analyzed with dose–response sigmoidal fit function to estimate minimum effective concentration and EC50 values. Chapter 2 122 2.9. Interaction of PARG NCs with the Caco-2 cells monolayer Entry of nanoparticles into Caco-2 cells was studied quantitatively by flow cytometry and qualitatively by confocal laser scanning microscopy (CLSM), for which DiD (λem= 644 nm) loaded nanoparticles were employed. For the flow cytometry study, Caco-2 cells were seeded in 24-well cell culture plates at a density of 5x105 cells per well and allowed to adhere for 48h until confluency. Cells were co-incubated with 400 μL of a DiD loaded nanoparticles suspension in HBSS (0.371 mg/mL NCs). After 2h of incubation with fluorescent NCs, cells were washed three times with PBS and detached from the plates by trypsinization. Cells were then centrifuged at 1500g, the supernatant was discarded, the cells were resuspended in PBS and fluorescence was measured using a BD FACSVerseTM flow cytometer (Becton Dickinson Biosciences, San Jose, CA, US). Cell fluorescence was quantified by measuring the fluorescence of DiD. For cell viability measurements, the propidium iodide reagent was employed. The reagent was added to each sample at a final concentration of 10µg/mL, and, after 10min of incubation, the fluorescence corresponding to dead cells was measured at 620nm (FL2). For each sample, 10,000 events were collected. The data were subsequently analyzed using the FlowJo data analysis software package (TreeStar, USA). For the CLSM study, the Transwell® inserts fixed in PFA 4% were gently washed in HBSS. Actin was stained with 200µL of alexa-phalloidine (1:50) in buffered HBSS+0.2% (v/v) Triton X-100 for 10 min in the dark to reveal cell borders, as described by des Rieux et al (32). Subsequently, inserts were washed in HBSS, cut and mounted on glass slides. Images were captured using a Zeiss™ confocal microscope (LSM 150). Data were analyzed by the Axio Vision software (vs 4.8) to obtain y-z, x-z and x-y views of the cells monolayers. 2.10. Toxicity and permeability on human intestinal tissues Concerning the toxicity on human intestinal tissue, jejunal tissue samples were collected from patients undergoing laparoscopic Roux-en-Y gastric bypass. Patients Chapter 2 123 had given full informed consent. The study has been reviewed and approved by the regional ethical review board. Tissue samples were immediately transferred into a vessel containing cold, oxygenated Krebs-Ringer buffer and quickly transported to the laboratory. Arriving the epithelium was dissected away from sub-epithelial tissues and mounted in horizontal as well as vertically oriented Ussing chambers with 9 mm openings between the two chambers. The chambers were kept at 37ºC and bubbled with 95% O2/5% CO2 for the duration of the experiment. Electrophysiology of the tissue was monitored throughout the experiment to assure continued tissue viability. After mounting, the tissues were allowed to equilibrate for 40 minutes with two medium exchanges (35, 36). NCs (1mg/mL) were then added to the chambers. At the end of the experiment, continued viability of the tissues was tested by addition of the cAMP-agonist forskolin. Viable tissue with oxidative metabolism will form cAMP in response to forskolin leading to an opening of CFTR Cl channels, the response was monitored as changes in potential difference and short-circuit current over the epithelium. In respect of the NC transport on intestinal tissue, studies has been performed on human jejunal tissues using Ussing chamber models as described above. DiD labeled PARG NCs up to 3.5mg/mL were added to the chambers, and samples from donor and acceptor side chambers were taken at regular intervals for the 120 min duration of the experiment. Permeability of fluorescently labeled NCs were analyzed in a plate reader (33). 2.11. Measurement of the trans-epithelial resistance (TEER) and insulin transport across the Caco-2 monolayer Caco-2 cell monolayers were cultured on tissue-cultured-treated PET filters (1 µm diameter, 1.1 cm2, Millipore Transwell 12 well/plate) and were used for experiments 21 days after seeding. The evaluation of PARG NCs was performed in Caco-2 cell monolayers with two NC concentrations: 0.5 mg/mL and 1 mg/mL. The variation in the TEER values for the cell monolayer integrity assessment was measured with a Chapter 2 124 Millicell-Electrical Resistance System (Endohm-12, Millipore Corp). Monolayers with a TEER values in the range of 800-1500 Ω cm2 were used. Simultaneously, samples were collected (500 µL) from the receiver compartment and the apical compartment 2h after NC cell monolayer exposure and the insulin concentrations were measured using a LC/MS. Cell monolayers were gently washed with NaCl 0.9% and frozen at -80°C for insulin quantification within the cells. Liquid chromatography (Shimadzu HPLC system LC 20AD) with a 150 x 2.1 mm - 5µm - 300Å HPLC C8 column (Interchim) was used for elution of insulin. The mobile phase A/B, where solvent A was H2O containing 0.1% formic acid and solvent B was acetonitrile containing 0.1% formic acid; the flow rate was 0.6mL/min to avoid pressure rise. 100 μL of tested sample was treated with 200 μL of chloroform / methanol / water at 1/ 1/ 0.3 and 100 μL of 0.1M NaOH, and then 40 μL of analyte was injected onto the column placed in an oven at 60 °C. The total run time was 13min. Detection was done by tandem mass spectrometry (Quantum ultra) in positive electrospray mode. System control and data processing were carried out using MassLynx software version 4.1. Spray voltage was 3.0kV, and sheath and auxiliary gas pressures were 50 and 15 (arbitrary units), respectively. The in-source CID energy was fixed at 12V, and capillary temperature was 350°C. Tube lens and collision energy values were optimized for insulin. Multiple reaction monitoring was used for the detection of the ion transitions. The multiple reaction monitoring transitions for analytes were as follows: m/z insulin/hexameric 709.805 > 731.76, m/z bovine insulin 1284.73 > 1104.60. Analytes were quantified by means of calibration curves using bovine insulin as internal standard. The standard curves showed linearity for creatine over a range of 0.025 - 10 μg·mL−1 for insulin. The methodology for this assay involves reduction with dithiolthreitol 45mM and alkylation with 100mM of iodoacetamide 100mM of intact insulin for measurement of the free B chain. Chapter 2 125 2.12. In vivo fluorescence imaging of DiD-loaded PARG NCs Two BALB/c mice were placed on a low manganese diet to reduce autofluorescence from normal mouse chow, and abdominal fur was removed by depilation where requested. One week later, 200μL of the DiD loaded PARG NCs (NC 20 mg/mL, DiD 10 μg/mL) were administered to mice by oral gavage. In vivo biodistribution was performed by total body scanning at different time points (0, 1, 3, 6, 24 hours) on isoflurane/oxygen anesthetized animals, using the MX2 scanner (ART, Montreal, Canada). 2.13. Bioactivity study of encapsulated insulin All animal experiments were reviewed and approved by the ethics committee of the University of Santiago de Compostela (procedures Prof. Carlos diéguez, 1500AE / 12 / FUN01 / FIS02 / CDG3) of according to the European and Spanish regulations for the use of animals in animal studies; performed therefore in compliance with the Directive 2010/63 / EU of the European Parliament and Council of 22nd September 2010 on the protection of animals used for scientific purposes; Spain Royal Decree 1201/2005, of October 10th, on the protection of animals used for experimental and other scientific purposes and under the Royal Decree 296/2008 of Spain 30th December on the protection of animals used for experimental and other scientific purposes, including teaching. Male Sprague-Dawley rats (247-272g) were obtained from from the Central Animals House of the University of Santiago de Compostela (Spain). The animals were fasted for 4h prior to experiments, with free access to water, and kept conscious during the whole experiment. A dose of insulin loaded PARG NCs (1 IU/kg) in a volume–weight ratio of 250μL:250g was administered subcutaneously (n=8). As control, plain insulin solution was administered to the animals following the same procedure at the same dosage (n=8). Blood samples were collected from the tail vein 30min prior to the subcutaneous administration to establish the baseline blood glucose level. At time point of 0.5, 1, 1.5, 2, 3, 4, 5, 6, 7 and 8 hours after Chapter 2 126 administration the blood samples were collected to monitor the glucose level change following the PARG NC or insulin administration. The glucose level was measured using a glucometer (GlucocardTM G+ meter, Arkray Factory, Japan). 2.14. Intra-duodenal / jejunal administration Male Sprague-Dawley rats (240-290g) were obtained from the Central Animals House of the University of Santiago de Compostela (Spain). The animals were fasted for 4h prior to experiments, with free access to water, and kept conscious during the whole experiment. PARG NCs were administered intra-duodenally or intro-jejunally to the rats at insulin dosage 50IU/kg body weight in a volume of 0.3mL through an intra-duodenal cannula operated 1 week before the experiment (n=8). As control, blank PARG NCs without insulin loading was administered to the animals following the same procedure (n=4). Additionally, an insulin solution in saline was subcutaneously injected at dose of 1 IU/kg body weight to a different group (n=8). Blood samples were collected from the tail vein 30min prior to the oral administration to establish the baseline blood glucose level. At time point of 0.5, 1, 1.5, 2, 3, 4, 5, 6, 7, 8h after dosing, the blood samples were collected to monitor the glucose level change following the PARG NC or insulin administration. The glucose level was measured by glucometer. 3.Results and discussion As presented in the introduction, the main objective of this work was to design, develop and characterize a peptide nanocarrier with a potential to confront all the barriers associated to the oral modality of administration. Because of our previous experience and positive results obtained with chitosan NCs (7, 8), we decided to adopt this NCs technology and to engineer it in order to obtain a NC prototype fulfilling the requirements associated to the oral administration. The tailored properties were i) a Chapter 2 127 capacity to load insulin, ii) a nanometric size and a monodispersed population, iii) a neutral / negative zeta potential and a capacity to interact with the intestinal epithelium and facilitate the transport of the associated peptide. Figure 1. Structural illustration of PARG NC with the compositions. Figure 2. Transmission electron micrographs of the PARG NCs 3.1. Physicochemical characteristics of NCs and association efficiency of insulin The schematic representation highlighting the components and the organization within the NCs is shown in Figure 1. The selection of these components and well as the appropriate concentrations of them were selected upon a thoughtful screening. Among the oils explored, oleic acid and Miglyol® 812N, the most commonly used one to formulate NCs (8, 9, 34-37), we selected oleic acid, because of the penetration enhancing effects of these long fatty acid chains as compared to the medium chains of Miglyol® 812N (38, 39). In addition, the important negative charge of oleic acid was supposed to facilitate the subsequent envelopment with cationic PARG. On the other hand, Span®80 was selected as surfactant among Span®80, Tween®80 and Labrasol®, Chapter 2 128 because it facilitate the attachment of PARG onto the NCs (confirmed by increased surface charge). The explanation can be that the PEG moieties in the other two emulsifiers saturated the NC surface area, which tremendously reduced the space for the interaction of PARG. Finally, SDC was selected as a co-surfactant because in addition to help reducing the NC particle size, it is known to form hydrophobic ionic pairs with insulin (21, 40), thereby facilitating its encapsulation. The selection of the prototypes resulting from the above-indicated screening process was based on the evaluation of specific properties. With regard to the particle size distribution and zeta potential, it was found that using specific amounts of the ingredients illustrated in Figure 1, it was possible to obtain PARG NCs with a hydrodynamic mean size of 178 ± 20 nm, low PDI (0.11) and a negative Z-Potential of -23 ± 2 mV. The low negative zeta potential was attributed to a compensation of the positive charge of PARG with the negative charges associated to other ingredients, such as SDC and oleic acid, as well as to the shielding effect of polyethylene oxide-polypropylene oxide (poloxamer 188). This barrier was considered to be a positive feature because, firstly, it is expected to increase the stability of the NCs when interacting with intestinal lipids and enzymes (41), and secondly, it may also facilitate the diffusion of the NCs across the mucus. On the other hand, using TEM, it was found that the NCs have a size smaller than the one observed by DLS. This could be due to the shrinking of the NCs during the drying process. The images also showed that the NCs have a spherical shape (Figure 2). In theory, the incorporation of insulin into the NCs is expected to depend on its electrostatic and hydrophobic interactions with various components of the NCs. Such interactions are obviously dependent on the solubility and ionization of insulin, both related to its isoelectric point (IP). In order to expose insulin molecules to different pH values, we adjusted the pH of insulin aqueous phase to different values and this Chapter 2 129 had an impact on the final pH of the formulations. As shown in Figure 3, when the final pH of the formulation was close to the insulin IP (IP=5.49), it was possible to reach an AE close to 80%. However, at pH values far away from this IP (either higher or lower), the insulin AE was clearly reduced (see Figure 3). The increased insulin entrapment at the pH close to its IP could be attributed to the predominance of hydrophobic interaction between the hydrophobic domains of this peptide with the components of the NC core. Figure 3. Influence of pH of the PARG NC suspension on insulin association efficiency. Mean± S.D., n = 3. Given the importance of pH on the insulin encapsulation process, we explored the inclusion of a buffer as the aqueous phase in order to ensure a final formulation pH close to 5.4. Among the different buffer systems investigated, we found that the use of 20 mM acetate buffer led to the formation of NCs with a size of 185± 6 nm, a low PDI and an AE of 88±5% (Table 2, determined by both direct and indirect method described in methodology section). Taking into account that the production yield of the NCs is 74.21±0.26% (w/w), the final insulin loading was 1.49% (w/w). 3 4.5 5 5.5 6.5 8 0 20 40 60 80 100 Association efficiency (%) pH of PARG NC suspension Chapter 2 136 A B Figure 8. Visualization of the interaction of DiD-labeled PARG NCs (0.371 mg/mL) with the Caco-2 cells monolayer after a 2h incubation time (n=3). (A) Flow cytometry profile showing the peak of DiD internalized cells (light blue) as compared to the non-fluorescent control cells (dark blue). (B) CLSM images (y-z, x-y and x-z sections) of the cell membranes upon staining with Alexa-phalloidine (green); DiD-labelled NCs are shown in blue color. 3.6.Effect on the transepithelial electrical resistance (TEER) Some cationic polymers, among them chitosan (7, 15), are well-known for their capacity to open the TJs, thereby altering the trans-epithelial resistance and facilitating paracellular transport of drugs. A recent report has also claim this functionality for PARG (46). In this study, we evaluated the effect of PARG NCs on the TEER of the Caco-2 monolayer and used a PARG solution as a control. The results indicated that the TEER value of Caco-2 cells was not affected at PARG NC concentration of 0.5 mg/mL (data not shown). However, when the concentration was increased up to 1mg/mL, both, PARG NCs and the free PARG, induced a significant TEER decrease (15% and 38%, respectively), after a 2-hour incubation time (Figure 9). Interestingly, after removing PARG NCs, a significant TEER recovery was observed at 24 hour, thus evidencing the transient opening of the TJs. In contrast, this recovery of the TEER value was not observed upon exposure to the free polymer. Chapter 2 137 Overall, this study confirmed the capacity of PARG polymer to open the intercellular TJ (46), and showed that by the incorporation of this polymer to the shell of the PARG NCs it is possible to modulate the permeation enhancing capacity of the polymer. This might also suggest that the toxicity observed on the Caco-2 cells maybe transitory. Figure 9. TEER assay on Caco-2 cell monolayers exposed to PARG NCs (1 mg/mL, containing 0.045mg/mL PARG and 0.014mg/mL insulin) or to PARG polymer (0.045 mg/mL) + insulin (0.014 mg/mL). 2h after exposure; 24h after removal of the prototypes from the cells. Data expressed as mean ± SD, n=3. Changes were considered statistically significant at P < 0.05: *p < 0.05 compared to the control group; ***p < 0.05 compared to PARG NC group. 3.7. Capacity of PARG NCs to enhance the transport of insulin As shown in previous sections, PARG NCs are able to adhere to the Caco-2 cell monolayer and alter in a transient manner the TEER. In a subsequent study, we studied whether or not these mechanistic details were translated into an enhanced insulin transport. The results showed that at upon a 2-hour incubation time, 3.54 ± 0.27% of insulin associated to the NCs was transferred to the basolateral compartment. This transport was significantly higher than the one observed for the physical mixture of PARG and insulin (2.73±0.32%) (Figure 10A). In addition, the amount of insulin detected inside the enterocytes was higher (1.29±0.53%) when the peptide was Chapter 2 138 administered associated to the NCs, with respect to the administration of the physical PARG-insulin mixture (0.67±0.08%, Figure 10B). Based on the observed adherence of the NCs to the monolayer and the transient changes in the TEER, it could be speculated that the enhanced insulin transport occurs by the paracellular pathway. Although the small amount (1.29 %) of insulin internalized in the monolayer also suggests the possibility of an enhanced intracellular uptake, as previously reported for octaarginine (49), it could be concluded that the dominant mechanism of transport is the one taking advantage of the paracellular route. In addition, the higher insulin transport achieved with PARG NCs, as compared to that achieved with the free PARG, could be explained by the co-localization of the peptide in association with PARG and other permeation enhancers present in the formulation. In fact, the oleic acid and bile salts present in the NCs core are known to increase the fluidity of cell membrane, and enhance the membrane permeability (16, 18, 39, 50-52). Figure 10. Apical to basolateral transport of insulin across the Caco-2 cell monolayer under 37°C, after 2h incubation with insulin loaded PARG NCs (1mg/mL, containing 0.045mg/mL PARG and 0.014mg/mL insulin) or PARG polymer (0.045 mg/mL) + insulin (0.014 mg/mL). Data expressed as mean ± SD, n=3. Changes are considered statistically significant at p < 0.05, evaluated by ANOVA following Tuke’s multiple comparison post hoc test (SigmaPlot SyStat Software Inc, San Joes, CA). A B Chapter 2 139 3.8. In vivo fluorescence imaging of DiD-loaded PARG NCs In order to have an in vivo preliminary estimation of the interaction of PARG NCs with the intestinal tract, we traced fluorescent DiDlabelled PARG NCs after oral administration to mice (Figure 11). The bioluminescent image on the left (acquired by the Optix OptiviewTM, ART, Montreal, Canada) shows the biodistribution of PARG NCs or free DiD dye following oral gavage to mice, and the histogram on the right shows the mean of total photons emitted from the regions of interest (ROI, the whole rat body). The images suggest that both, the PARG NCs and the free DiD dye, remain associated to the gastro-intestinal tract for up to 24 hours (Figure 11). The association of the free amphiphilic dye to the mucosa could be explained by its affinity for the cell’s membrane. However, the analysis of the florescence intensity gave some preliminary evidence of the greater retention of the NCs as compared to the free dye. More detailed in vivo studies are needed in order to confirm this interaction of the NCs with the intestinal mucosa. Chapter 2 140 Figure 11: Fluorescence images (on the left) of representative mice at 0h, 1h, 3h, 6h and 24h following oral administration of (A) DiD labelled PARG NCs, (B) free DiD dye; and the mean of total photons ± SD emitted from regions of interest (ROI) around the GIT (on the right). 3.9. Bioactivity of encapsulated insulin and in vivo efficacy of PARG NCs It is well known in the biotechnology field that the formulation process may result in the inactivation of labile macromolecules, such as peptides (53). Within this context, and before studying the in vivo efficacy of the oral formulation, we analyzed the bioactivity of the peptide following subcutaneous (s.c.) administration. Thus, insulin-loaded PARG NCs were administered subcutaneously to fasted (4h) healthy rats, using an insulin saline solution as control. The blood glucose level was Chapter 2 141 normalized taking the 0h mean glucose baseline value as 100%. As shown in Figure 12, following the s.c. injection, both, the insulin solution and insulin-loaded NCs, exhibited a very similar profile, where a drastic decrease in the glucose level was observed in 0.5h and the normal levels were recovered in 3h. From these results, it can be concluded that the mild conditions formulation process adopted for this system has not affected to the bioactivity of insulin. Finally, the in vivo efficacy study was performed following either intra-duodenal or intra-jejunal administration (50IU/kg) to conscious healthy rats after 4h fasting (Figure 12). In both cases, the administration of insulin-loaded PARG NCs led to a slight blood glucose level decrease along the first 3 hours of the assay in comparison to blank PARG NCs controls. However, there were no significant differences in the responses observed for insulin-loaded NCs and the blank NCs. Overall, it could be concluded that the amount of insulin that reached blood circulation, after intra-duodenal or intra-jejunal administration (50 IU/kg), was much lower than the one achieved after sc. administration of 1 IU/kg insulin. Figure 12. Standardized hypoglycemic effect in healthy rats following subcutaneous administration of insulin-loaded PARG NCs and insulin saline solution at 1 IU/kg, intra-duodenal administration of insulin-loaded PARG NCs at 50 IU/kg, intra-jejunal 0 1 2 3 4 5 6 7 8 9 0 20 40 60 80 100 120 Blood glucose level (%) Time (h) s.c.insulin s.c. NC Intra-duodenal NC Intra-jejunal NC Placebo Chapter 2 142 administration of insulin-loaded PARG NCs at 50 IU/kg, and intra-jejunal administrated blank PARG NC as placebo. Data represents the mean ± S.E., n=8 for all the groups except for placebo (n=4). This poor in vivo performance is somehow contradictory with the rational design of PARG NCs. In fact, the observed preservation of the insulin activity during the formulation process, the adequate stability in the intestinal media and the remarkable capacity of this formulation to promote the transport of insulin across a Caco-2 cells monolayer (Figure 9) would rather suggest a potential for this formulation to facilitate insulin oral absorption. A number of hypothesis have been formulated to explain the limited performance of the nanocarriers. First, the viscosity of the NC formulation, which might prevent the adequate mixing with the mucus fluids; second, despite of the controlled release properties of the formulation, there is the possibility that some enzyme (pancreatin) molecules may interact with the NCs and promote insulin degradation; third, although NCs may interact with the intestinal epithelium, it is possible that the interaction is insufficient in the in vivo situation; finally, even if the NCs interact with the epithelium, the associated insulin maybe retained and even degraded at the intracellular level. Finally, it is also possible that the in vitro assays and/ or the in vivo experimental conditions used in the reported studies have a limited predictive value of the performance of these formulations. Specific studies, i.e. mucodiffusion and quantification of insulin absorption are underway to validate these hypotheses. From the studies performed so far we could conclude that the rational development of oral peptide delivery formulations and the translation of the in vitro data into the in vivo situation possess significant difficulties. However, the information reported here is supposed to help identifying the critical steps to be considered in the design of such formulations. Further work of PARG formulations aimed at enhancing the peptide Chapter 2 143 loading, as well as the incorporation of the NCs in a final dosage form (beads or capsules) might additionally help increase the performance of the formulation. 4. Conclusion In this work we report the rational design of NCs consisting of an oily core made of penetration enhancers (oleic acid and SDC) and a polymer shell made of PARG, a polymer that is known for its ability to interact and increase the permeability of the cell membranes. 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