Nanotechnology for Oral Drug Delivery: From Concept to Applications [1 ed.] 0128180382, 9780128180389

Table of contents :
Cover
NANOTECHNOLOGY
FOR ORAL DRUG
DELIVERY
From Concept to Applications
Copyright
Contributors
Foreword
Part I: Biological aspects and properties of nanomaterials for oral drug delivery
Oral drug delivery: Overview
Introduction
References
Organization of the intestinal mucosa and barriers to oral drug delivery
Introduction
Structure of the small intestine
Cellular compositions
Tight junctions or zonula occludens
Adherens junctions or zonula adherens
Desmosomes or macula adherens
Metabolizing enzymes
Efflux pumps
Transporters for transcytosis
Mechanism of crossing the intestinal mucosal barrier
Passive diffusion through transcellular pathways
Passive diffusion through paracellular pathways
Models to study intestinal absorption
In vitro models
In vivo models
Conclusions
References
Nanomaterials for oral drug administration
Introduction
The rational design of nanocarriers for oral drug formulation
Types of nanomaterials
Nanocrystals and nanosuspensions
Polymeric-based nanocarriers
Nanospheres and nanocapsules
Polymeric micelles
Nanogels
Dendrimers
Polymer-drug conjugates
Lipid-based nanocarriers
Liposomes
Nanoemulsions
Solid lipid nanoparticles (SLN) and nanostructured lipid carriers (NLC)
Inorganic based-nanocarriers
Carbon nanotubes
Nonporous nanomaterials
Mesoporous nanomaterials
Emerging trends in oral nanoformulation: Hybrid, protein and stimuli-responsive nanocarriers
Hybrid nanocarriers
Protein nanocarriers
Stimuli-responsive nanocarriers
Nanomaterials for the targeting of specific GIT regions
Targeting the stomach
Targeting the small intestine
Colon targeting
Conclusions and future perspectives
References
Mucoadhesive and mucus-penetrating polymers for drug delivery
Introduction
Definitions and fundamentals
Bio- and mucoadhesion
Oral mucosa as route for drug administration
Reasons for developing oral mucosal drug delivery systems
Desired physicochemical and biopharmaceutical characteristics of drugs
Desired physicochemical characteristics of materials
Mucoadhesive and mucus-penetrating materials
Mucoadhesive polymers
Natural mucoadhesive polymers
Semisynthetic mucoadhesive polymers
Synthetic mucoadhesive polymers
New generation of mucoadhesive polymers
Mucus-penetrating polymers
Oral mucosal drug delivery systems
Strategies
Technology
Types of systems or pharmaceutical dosage forms
Safety and toxicity
In-vitro evaluation
In-vivo evaluation
Regulatory status
Patenting and market
Conclusions and future perspectives
References
Size, shape and surface charge considerations of orally delivered nanomedicines
Introduction to nanomedicine for oral drug delivery
Challenges associated with oral nDDS
Size, shape and surface chemistry considerations for orally delivered nanomedicine
Particle size and shape
Surface properties of nanoparticles
Drug loading and release profile
Summary
References
Modified drug release: Current strategies and novel technologies for oral drug delivery
Part 1: Evolution of oral modified release dosage forms
Part 2: Current oral formulation strategies for modified release
Extended release formulations
Dissolution-controlled formulations
Diffusion-controlled formulations
Osmosis-based formulations
Ion exchange-based formulations
Delayed release formulations
Targeted release formulations
Gastroretentive (GR) devices
Colon-targeted drug delivery systems
Novel technologies for modified drug release
Microneedle pills for oral drug delivery
Three-dimensional (3D) printing medicines
Conclusions and future perspectives
References
Delivery platforms for oral drug administration
Introduction
Delivery platforms for oral administration of (bio)pharmaceuticals
Effect of the size and shape of the drug delivery system
Effect of the surface of the drug delivery system
Effect of the composition and the preparation method of the drug delivery system
Targeted drug delivery systems
Conclusions
References
Further reading
(Trans)buccal drug delivery
Introduction
Oral cavity: Anatomic and physiologic features
Buccal mucosa as barrier for drug penetration/permeation
Strategies to target the buccal mucosa
Penetration enhancers
Mucoadhesion
Enzyme inhibitors
The advantages of nanoparticles for buccal delivery
Polymeric nanoparticles
Polymeric micelles
Lipid-based nanoparticles
Liposomes
Dosage forms for the buccal delivery of nanoparticles
Films
Solid matrices
Gels
Final remarks
References
Part II: Advanced technologies for oral delivery applications
Spray-drying for the formulation of oral drug delivery systems
Introduction
Spray-drying
Principle, equipment configurations, main advantages and challenges
Parameters and variables affecting particle formulation
Factors affecting particle formation mechanisms
Feed solution properties
Process parameters
Scale-up considerations
Formulation of drugs/drug delivery systems for oral administration using the spray-drying technology
Conventional methods versus spray-drying of micro- and nano-particle formulations
Common excipients used in the preparation of drug formulations prepared by spray-drying
Preparation of particulate systems with spray-drying and applications in oral drug delivery
Microparticles
Nanoparticles
Pure drug particles
Conclusions and future perspectives
References
Microdevices to successfully deliver orally administered drugs
Introduction
Design of microdevices
Materials for the fabrication of microdevices
Non-biodegradable materials
Biodegradable materials
Methods for microfabrication of oral drug delivery devices
Lithography
Photolithography
Soft lithography
Imprint lithography
Embossing and punching
Additive manufacturing
Other methods
Loading techniques for drug formulations
Loading of liquids
Inkjet printing
Loading of hydrogels
Loading of powders into the microdevices
Manual loading of the microdevices
Hot embossing
CO2 impregnation
Lid formation on the cavity of microdevices
pH-sensitive lids
Mucoadhesive lid formation
In vitro and ex vivo studies
Characterization techniques of microdevices and their loading
Bioadhesion of microdevices
In vitro drug release from microdevices
In vitro drug transport and toxicity
In vivo testing and applications
Future perspectives
Conclusion
References
Batch and microfluidic reactors in the synthesis of enteric drug carriers
Introduction
Nanotechnology in oral drug delivery
Nanotechnology approaches for drug delivery carriers to cross the intestinal lining
Nanotechnology approaches for attaching drug delivery carriers to the intestinal mucosa
Delivery to the systemic circulation-Passage through the intestinal barrier
pH-dependent and pH-independent targeted delivery to specific sites of the intestine
pH-dependent drug delivery
Copolymers of methyl acrylate, methyl methacrylate and methacrylic acid
Polymers based on cellulose derivatives
Polymers based on polyvinyl derivatives
pH-independent drug delivery
Microfluidics for the synthesis of micro and nanocarriers in oral delivery
Conclusions
References
3D printing in oral drug delivery
Oral delivery
Strategies to overcome physiological limitations of the gastrointestinal (GI) tract
Introduction to the different release mechanisms
Fixed dose combination drug therapy (FDCDT)
Personalized medicine
Motivation for the patient-tailored medicine
Prerequisites for the production of personalized medicine
Where do we need the 3D printing?
Patient perception and preferences of the medicine regarding shape, color, embossing, flavor, and acceptability of p ...
Introduction to additive manufacturing
Fused deposition modeling
Semi-solid extrusion
Powder-bed printing
Regulatory challenges
The first 3D printed drug product on the market
Requirements for raw materials, printers and manufacturing procedures
Supply chain
Data management from the regulatory perspective
Future perspective
Robotic devices
Digital technology and Internet of Things: Fabrication, identification, anti-counterfeiting
Personalized medicine for veterinary purposes
Concluding remarks
References
Further reading
Part III: Methods for the evaluation of oral drug delivery systems
3D intestinal models towards a more realistic permeability screening
Introduction
3D intestinal models
Multilayered models
Organoids
Gut-on-a-chip models
Limitations and future perspectives
Conclusion
References
In vitro relevant information for the assessment of nanoparticles for oral drug administration
General considerations of oral drug delivery
Barriers to oral drug delivery
Gastrointestinal juices and microbiota
Mucus
Cellular barriers
Understanding the biological-material interface
Electrolytes and pH
Enzymes and active surface molecules
Mucus
Intestinal epithelium
Nanoparticles for oral drug delivery: In vitro characterization techniques
Stability in gastrointestinal fluids
Mucoadhesion and mucodiffusion
Handmade techniques
Analysis of muco-interaction
Analysis of mucodiffusion
Specialized-equipment techniques
Interaction with the intestinal epithelium
Conventional 2D cell culture
Enterocyte-like model
Mucus enterocyte model
Follicle associated epithelium model
3D cell culture
Engineered intestinal tissues
Microfluidic-based approaches
Conclusions and future perspectives
References
In vivo testing of orally delivered nanoparticles
Introduction
Selection of animal models
Administration technique
Oral gavage
Intragastric infusion
Intestinal instillation
Syringe-feeding technique
Experimental techniques
Pharmacodynamics study
Pharmacokinetic profiling
Organ/tissue analysis
Imaging systems
Radioactive labeling
Optical imaging
Histological analysis
In vitro-In vivo correlation
Ethical considerations
Conclusion
References
Part IV: Pharmaceutical industry perspective
Industrial perspectives and future of oral drug delivery
Introduction: Oral nano-drug delivery
Current market status
Challenges
Gap analyses from benchtop to good manufacturing practice (GMP) production
Understanding bench top methods and their limitations for up scaling
Top-down methods
Bottom-up methods
Scale-up of nanoparticle production methods
Scale-up of nanoparticle downstream methods
Regulatory aspects: Regulations and guidelines
Conclusions and future perspectives
Annex I. Existing guidance documents expected to be applied for nano-enabled pharmaceutical products
ICH guidelines
Quality
Safety
FDA guidance for industryeehttps://www.fda.gov/animal-veterinary/guidance-regulations/guidance-industry.
FDA nanotechnology guidance documentsffhttps://www.fda.gov/science-research/nanotechnology-programs-fda/nanotechno
References
Index
A
B
C
D
E
F
G
H
I
K
L
M
N
O
P
Q
R
S
T
U
V
W
X
Z
Back Cover

Citation preview

NANOTECHNOLOGY FOR ORAL DRUG DELIVERY

NANOTECHNOLOGY FOR ORAL DRUG DELIVERY From Concept to Applications Edited by

JOÃO PEDRO MARTINS Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland

 HELDER A. SANTOS Full Professor, Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland Full Professor, Helsinki Institute of Life Science (HiLIFE), University of Helsinki, Helsinki, Finland

Academic Press is an imprint of Elsevier 125 London Wall, London EC2Y 5AS, United Kingdom 525 B Street, Suite 1650, San Diego, CA 92101, United States 50 Hampshire Street, 5th Floor, Cambridge, MA 02139, United States The Boulevard, Langford Lane, Kidlington, Oxford OX5 1GB, United Kingdom © 2020 Elsevier Inc. All rights reserved. No part of this publication may be reproduced or transmitted in any form or by any means, electronic or mechanical, including photocopying, recording, or any information storage and retrieval system, without permission in writing from the publisher. Details on how to seek permission, further information about the Publisher’s permissions policies and our arrangements with organizations such as the Copyright Clearance Center and the Copyright Licensing Agency, can be found at our website: www.elsevier.com/permissions. This book and the individual contributions contained in it are protected under copyright by the Publisher (other than as may be noted herein). Notices Knowledge and best practice in this field are constantly changing. As new research and experience broaden our understanding, changes in research methods, professional practices, or medical treatment may become necessary. Practitioners and researchers must always rely on their own experience and knowledge in evaluating and using any information, methods, compounds, or experiments described herein. In using such information or methods they should be mindful of their own safety and the safety of others, including parties for whom they have a professional responsibility. To the fullest extent of the law, neither the Publisher nor the authors, contributors, or editors, assume any liability for any injury and/or damage to persons or property as a matter of products liability, negligence or otherwise, or from any use or operation of any methods, products, instructions, or ideas contained in the material herein. Library of Congress Cataloging-in-Publication Data A catalog record for this book is available from the Library of Congress British Library Cataloguing-in-Publication Data A catalogue record for this book is available from the British Library ISBN : 978-0-12-818038-9 For information on all Academic Press publications visit our website at https://www.elsevier.com/books-and-journals

Publisher: Wolff, Andre Acquisitions Editor: Hill-Parks, Erin Editorial Project Manager: Young, Sam W Production Project Manager: Raviraj, Selvaraj Cover Designer: Pearson, Victoria Typeset by SPi Global, India

Contributors

Beatrice Albertini Department of Pharmacy and BioTechnology, University of Bologna, Bologna, Italy Andreia Almeida i3S—Instituto de Investigac¸a˜o e Inovac¸a˜o em Sau´de; Instituto de Ci^encias Biomedicas Abel Salazar, Universidade do Porto, Porto, Portugal Marı´a Jose Alonso Center for Research in Molecular and Chronic Diseases (CIMUS); Department of Pharmaceutics and Pharmaceutical Technology, School of Pharmacy, Campus Vida, University of Santiago de Compostela, Santiago de Compostela, Spain Manuel Arruebo Department of Chemical Engineering, Aragon Institute of Nanoscience (INA), University of Zaragoza and Instituto de Ciencia de Materiales de Arago´n (ICMA), Universidad de Zaragoza-CSIC, Zaragoza; Networking Research Center on Bioengineering, Biomaterials and Nanomedicine, CIBER-BBN, Madrid, Spain Cla´udia Azevedo i3S—Instituto de Investigac¸a˜o e Inovac¸a˜o em Sau´de; Instituto de Ci^encias Biomedicas Abel Salazar, Universidade do Porto, Porto, Portugal Abdul W. Basit UCL School of Pharmacy, University College London, London, United Kingdom Serena Bertoni Department of Pharmacy and BioTechnology, University of Bologna, Bologna, Italy Johan Peter Boetker Department of Pharmacy, Faculty of Health and Medical Sciences, University of Copenhagen, Copenhagen, Denmark Maria Cristina Bonferoni Department of Drug Sciences, University of Pavia, Pavia, Italy Marcos Luciano Bruschi Laboratory of Research and Development of Drug Delivery Systems, Department of Pharmacy, State University of Maringa, Maringa, Parana, Brazil Jessica Bassi da Silva Laboratory of Research and Development of Drug Delivery Systems, Department of Pharmacy, State University of Maringa, Maringa, Parana, Brazil Henry P. Diehl Department of Pharmaceutical Chemistry, The University of Kansas, Lawrence, KS, United States Franca Ferrari Department of Drug Sciences, University of Pavia, Pavia, Italy

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Contributors

Mo´nica P. A. Ferreira Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland Natalja Genina Department of Pharmacy, Faculty of Health and Medical Sciences, University of Copenhagen, Copenhagen, Denmark Nazende G€ unday T€ ureli € MyBiotech GmbH, Uberherrn, Germany Jouni Hirvonen Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland Tushar Kumeria School of Pharmacy, The University of Queensland; Translational Research Institute, Mater Research Institute-The University of Queensland, Brisbane, QLD, Australia Robert Langer Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, MA, United States Marı´a Victoria Lozano Cellular Neurobiology and Molecular Chemistry of the Central Nervous System Group, Faculty of Pharmacy; Regional Centre of Biomedical Research (CRIB), University of Castilla-La Mancha (UCLM), Albacete, Spain Maria Helena Macedo i3S—Instituto de Investigac¸a˜o e Inovac¸a˜o em Sau´de; Instituto de Ci^encias Biomedicas Abel Salazar, Universidade do Porto, Porto, Portugal Joa˜o Pedro Martins Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland Chiara Mazzoni Department of Health Technology, Technical University of Denmark, Lyngby, Denmark Line Hagner Nielsen Department of Health Technology, Technical University of Denmark, Lyngby, Denmark Nadia Passerini Department of Pharmacy and BioTechnology, University of Bologna, Bologna, Italy Amirali Popat School of Pharmacy, The University of Queensland; Translational Research Institute, Mater Research Institute-The University of Queensland, Brisbane, QLD, Australia Vivitri D. Prasasty Faculty of Biotechnology, Atma Jaya Catholic University of Indonesia, Jakarta, Indonesia Veronique Preat Advanced Drug Delivery and Biomaterials, Louvain Drug Research Institute, Universite catholique de Louvain, Brussels, Belgium

Contributors

Jukka Rantanen Department of Pharmacy, Faculty of Health and Medical Sciences, University of Copenhagen, Copenhagen, Denmark Prarthana Rewatkar School of Pharmacy, The University of Queensland, Brisbane, QLD, Australia Silvia Rossi Department of Drug Sciences, University of Pavia, Pavia, Italy Marco Ruggeri Department of Drug Sciences, University of Pavia, Pavia, Italy Giuseppina Sandri Department of Drug Sciences, University of Pavia, Pavia, Italy Manuel J. Santander-Ortega Cellular Neurobiology and Molecular Chemistry of the Central Nervous System Group, Faculty of Pharmacy; Regional Centre of Biomedical Research (CRIB), University of Castilla-La Mancha (UCLM), Albacete, Spain Helder A. Santos Full Professor, Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy; Full Professor, Helsinki Institute of Life Science (HiLIFE), University of Helsinki, Helsinki, Finland Bruno Sarmento i3S—Instituto de Investigac¸a˜o e Inovac¸a˜o em Sau´de, Universidade do Porto, Porto; CESPU, Instituto de Investigac¸a˜o e Formac¸a˜o Avanc¸ada em Ci^encias e Tecnologias da Sau´de, Gandra, Portugal Victor Sebastian Department of Chemical Engineering, Aragon Institute of Nanoscience (INA), University of Zaragoza and Instituto de Ciencia de Materiales de Arago´n (ICMA), Universidad de Zaragoza-CSIC, Zaragoza; Networking Research Center on Bioengineering, Biomaterials and Nanomedicine, CIBER-BBN, Madrid, Spain Katia P. Seremeta Department of Basic and Applied Sciences, National University of the Chaco Austral; National Scientific and Technical Research Council (CONICET), Pcia. Roque Sa´enz Pen˜a, Chaco, Argentina Neha Shrestha Advanced Drug Delivery and Biomaterials, Louvain Drug Research Institute, Universite catholique de Louvain, Brussels, Belgium Teruna J. Siahaan Department of Pharmaceutical Chemistry, The University of Kansas, Lawrence, KS, United States Alejandro Sosnik Laboratory of Pharmaceutical Nanomaterials Science, Department of Materials Science and Engineering, Technion-Israel Institute of Technology, Haifa, Israel

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Contributors

Sabrina Barbosa de Souza Ferreira Laboratory of Research and Development of Drug Delivery Systems, Department of Pharmacy, State University of Maringa, Maringa, Parana, Brazil Sarah J. Trenfield UCL School of Pharmacy, University College London, London, United Kingdom Akif Emre T€ ureli € MyBiotech GmbH, Uberherrn, Germany Barbara Vigani Department of Drug Sciences, University of Pavia, Pavia, Italy Aldyn Wildey Department of Pharmaceutical Chemistry, The University of Kansas, Lawrence, KS, United States

Foreword

This book addresses two of the most significant areas of drug delivery today—oral drug delivery and nanotechnology. The book boasts a number of excellent scientists as authors from around the world. It covers chemical, biological, and engineering aspects of oral drug delivery using micro and nano-particles. In particular, the book contains a number of chapters on the biology and barriers to oral drug delivery, as well as ways to study them. These chapters include an examination of the intestinal mucosa as a barrier to drug delivery as well as the development of threedimensional intestinal models and how to use these models to realistically screen for drug permeability. A number of chapters focus on biomaterials. There are chapters on nanomaterials for oral drug administration, and on microadhesive and mucus penetrating particles. Other chapters discuss manufacturing considerations. For example, three-dimensional printing and spray drying are examined in detail, as are the use of different reactions in the synthesis of enteric drug carriers. Microdevices for oral drug delivery are also discussed. In addition, an industrial perspective is provided. The book has chapters examining a number of other important topics. These include chapters on the size and shape of particulate carriers, on targeting such systems to increase transport of orally administered drugs, and on other approaches for modifying drug release. Buccal drug delivery is also discussed. Finally, both in vitro and in vivo approaches for examining oral delivery of nanoparticles are provided. This book should be a useful reference to scientists in universities and companies alike who are working on oral drug delivery and nanoparticulate systems. This is an area that is also expected to increase in importance in the future. Robert Langer MIT Department of Chemical Engineering, Massachusetts Institute of Technology, Cambridge, MA, United States

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CHAPTER 1

Oral drug delivery: Overview
lder A. Santosb,c João Pedro Martinsa, He a

Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland b Full Professor, Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland c Full Professor, Helsinki Institute of Life Science (HiLIFE), University of Helsinki, Helsinki, Finland

1. Introduction Nanomedicine consists on the medical application of nanotechnology towards the implementation of advanced diagnostic and therapeutic strategies that can improve the delivery of a wide range of therapeutics. Over the years, nanotechnology has been finding its application in all the specializations of drug delivery, driven by the rapid evolution of newly developed drug delivery strategies, technologies and treatment modalities [1]. The approval of new drugs and the possibility to use biologically active molecules that were previously considered as undevelopable due to their suboptimal pharmaceutical properties have also shed a light on the potential of nanotechnology to overcome persistent challenges of drug delivery. After several years of investigation, nanotechnology-based drug carrier systems have demonstrated to address the shortcomings of conventional therapies in multiple ways, allowing for targeted/site-specific and controlled drug delivery, with reduced side effects, and a more effective therapeutic outcome [2]. The advantages of using nano-engineered drug delivery systems over their non-formulated free drug counterparts is also supported by the increasing number of clinical trials and clinically approved nanotechnology-based drug delivery products in the market [1, 3]. However, of the different routes of administration, the intravenous drug delivery receives the most attention in both preclinical and clinical settings [3]. The intravenous administration of drugs brings together a set of drawbacks related to the necessity of frequent and painful injections, strict sterility conditions, high costs and, consequently, low patient compliance [4]. Therefore, it is of utmost importance to explore alternative routes of drug administration. Pulmonary [5], nasal [6], vaginal and rectal [7], and oral [8] modalities are often studied, with the latter being, unarguably, the preferred route of administration. Oral delivery is the most widely used form of drug administration due to ease of ingestion, cost effectiveness, and flexibility of the dosage regimen [9]. Neither sophisticated sterile

Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00001-6

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Nanotechnology for oral drug delivery

facilities, nor the direct involvement of health care professionals is required. Additionally, oral drug delivery systems are painless and have the highest patient compliance [10]. However, the path of orally administered therapeutics is not forthright. Low drug stability in the hostile conditions of the stomach, low drug solubility, the continuous mucus secretion and, ultimately, the presence of a tight intestinal epithelial barrier turn drug penetration and absorption into an intricate process, impairing the bioavailability of orally administered drugs [11]. Therefore, orally delivered therapeutics need to overcome a myriad of hurdles and still achieve therapeutic levels. The field of materials science has been providing advanced tools to develop robust drug carriers that are able to overcome the gastrointestinal hurdles that impair the bioavailability of therapeutics, and to deliver them in a site-specific and controlled manner. In this context, a great variety of biocompatible materials have been investigated in the form of carrier systems to overcome low drug solubility, dissolution, and poor bioavailability after oral administration [12, 13]. The exploitation of these materials is evolving alongside comprehensive studies on the influence of the size, shape and surface properties of these drug carriers in the success of oral drug administration [14, 15]. A wide range of fabrication methods and characterization techniques have also been explored to develop new and more efficient drug delivery systems, with the aim of accelerating their translation to the clinics [9]. Hence, the choice of materials, design aspects, and fabrication methods play a fundamental role in dictating the success of advanced drug delivery systems for oral administration. Additionally, the validation of their capacity to overcome the harsh conditions of the gastrointestinal tract and to enhance the drug transmucosal and transepithelial transport requires the use of suitable and clinically relevant models that can recapitulate the biological environments. For this reason, a great number of in vitro models have been established throughout the years to allow for an in-depth understanding of the phenomena occurring at the biological-material interface, and for carrying out systematic studies to assess the permeation coefficients of free drugs and drug-loaded carriers through the intestinal mucosal barrier [16]. Nonetheless, in vivo models remain crucial for validating in vitro experimental findings, and are still the gold standard method used during the preclinical development. In vivo models are thus essential to study the pharmacokinetics and pharmacodynamics, as well as the toxicity of nanotechnology-based drug delivery systems [17]. In spite of the extensive preclinical research efforts, the translation of drug nanocarrier systems into the clinic is severely retarded by a lack of standardization of the manufacturing and approval procedures, together with uncertainties in the evaluation of their quality and safety via traditional methods [18]. Therefore, it is urgent to establish cost effective and straightforward regulatory mechanisms that can foster the translation of these drug delivery systems from the bench to the bedside.

Oral drug delivery

The present book discusses the current challenges of oral drug delivery, broadly revising the different physicochemical barriers faced by orally administered drugs, and strategies to improve their intestinal permeability. A comprehensive overview of the most promising and up-to-date engineered and surface functionalized drug nanocarriers for oral drug delivery is presented. The relevance of controlling the physicochemical properties of the developed nanoformulations, from size and shape to the drug release profiles, are also reviewed. A variety of fabrication methods and characterization techniques for the development of novel and robust delivery platforms for the oral administration of therapeutics is described. Advances in both in vitro and in vivo scenarios are discussed, focusing on the interactions at the biological–material interface. Finally, the importance of the industry on the development of nanotechnology-based oral drug formulations in terms of manufacturing and commercialization is also reviewed. From essentials to applications, and motivated by the potential impact of nanomedicine in human healthcare, this book provides both the scientific and medical communities with a broad range of knowledge and expertise in the field of oral drug delivery with the use of nanotechnology.

References [1] Anselmo AC, Mitragotri S. Nanoparticles in the clinic: an update. Bioeng Transl Med 2019;4(3). [2] Wolfram J, Zhu M, Yang Y, Shen J, Gentile E, Paolino D, et al. Safety of nanoparticles in medicine. Curr Drug Targets 2015;16(14):1671–81. [3] Anselmo AC, Mitragotri S. Nanoparticles in the clinic. Bioeng Transl Med 2016;1(1):10–29. [4] Sastry SV, Nyshadham JR, Fix JA. Recent technological advances in oral drug delivery—a review. Pharm Sci Technol Today 2000;3(4):138–45. [5] Sung JC, Pulliam BL, Edwards DA. Nanoparticles for drug delivery to the lungs. Trends Biotechnol 2007;25(12):563–70. [6] Nochi T, Yuki Y, Takahashi H, Sawada S, Mejima M, Kohda T, et al. Nanogel antigenic proteindelivery system for adjuvant-free intranasal vaccines. Nat Mater 2010;9(7):572–8. [7] Mesquita L, Galante J, Nunes R, Sarmento B, das Neves J. Pharmaceutical vehicles for vaginal and rectal administration of anti-HIV microbicide nanosystems. Pharmaceutics 2019;11(3):145. [8] Reinholz J, Landfester K, Mail€ander V. The challenges of oral drug delivery via nanocarriers. Drug Deliv 2018;25(1):1694–705. [9] Date AA, Hanes J, Ensign LM. Nanoparticles for oral delivery: design, evaluation and state-of-the-art. J Control Release 2016;240:504–26. [10] Lamson NG, Berger A, Fein KC, Whitehead KA. Anionic nanoparticles enable the oral delivery of proteins by enhancing intestinal permeability. Nat Biomed Eng 2020;4:84–96. [11] Ensign LM, Cone R, Hanes J. Oral drug delivery with polymeric nanoparticles: the gastrointestinal mucus barriers. Adv Drug Deliv Rev 2012;64(6):557–70. [12] Chou LYT, Ming K, Chan WCW. Strategies for the intracellular delivery of nanoparticles. Chem Soc Rev 2011;40(1):233–45. [13] Boyd BJ, Bergstr€ om CAS, Vinarov Z, Kuentz M, Brouwers J, Augustijns P, et al. Successful oral delivery of poorly water-soluble drugs both depends on the intraluminal behavior of drugs and of appropriate advanced drug delivery systems. Eur J Pharm Sci 2019;137:104967. [14] Banerjee A, Qi J, Gogoi R, Wong J, Mitragotri S. Role of nanoparticle size, shape and surface chemistry in oral drug delivery. J Control Release 2016;238:176–85.

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[15] Yun Y, Cho YW, Park K. Nanoparticles for oral delivery: targeted nanoparticles with peptidic ligands for oral protein delivery. Adv Drug Deliv Rev 2013;65(6):822–32. [16] Arumugasaamy N, Navarro J, Kent Leach J, Kim PCW, Fisher JP. In vitro nodels for studying transport across epithelial tissue barriers. Ann Biomed Eng 2019;47(1):1–21. [17] Gamboa JM, Leong KW. In vitro and in vivo models for the study of oral delivery of nanoparticles. Adv Drug Deliv Rev 2013;65(6):800–10. [18] Desai N. Challenges in development of nanoparticle-based therapeutics. AAPS J 2012;14(2):282–95.

CHAPTER 2

Organization of the intestinal mucosa and barriers to oral drug delivery Henry P. Diehla, Aldyn Wildeya, Vivitri D. Prasastyb, Teruna J. Siahaana a Department of Pharmaceutical Chemistry, The University of Kansas, Lawrence, KS, United States Faculty of Biotechnology, Atma Jaya Catholic University of Indonesia, Jakarta, Indonesia

b

1. Introduction When developing drugs from the bench to the bedside, potential drug molecules must go through multiple steps in the process. One of these is the evaluation of the absorption properties of the potential drug using in vitro and in vivo models of intestinal mucosal barriers. The intestinal mucosa absorption of a potential drug is influenced in part by its physicochemical properties [1]. The selection of a delivery route is also critical for the clinical success of the drug candidate. Because of the high cost of drug development, many methods have been created to rapidly select a potential drug or “druggable” candidate. In addition, metabolism, pharmacokinetics (DMPK), and toxicity profiles need to be investigated. Oral delivery is the most convenient way to administer drugs to patients; therefore, during drug development stages, many drug candidates are evaluated for their ability to be delivered orally. Small drug molecules require the appropriate physicochemical properties for their absorption by the intestinal mucosal barrier. In this case, the molecules must pass through the mucosa layer followed by crossing the intestinal epithelial cell layer into the systemic circulation for their distribution to the target organ or tissue(s) [1]. The gastrointestinal (GI) tract has an effective location for food digestion as well as waste excretion. The resulting nutrients from digested foods are absorbed along the small intestine. At the same time, the GI tract protects the systemic circulation from hazardous pathogens and toxins from the lumen by acting as a selective barrier [2, 3]. Similar to nutrients, drug molecules are absorbed in the GI tract—especially the small intestine. Thus, understanding the structure of the small intestine along with its biochemical composition is necessary for developing drugs that will be delivered orally. In addition, knowing the physicochemical properties and physiological mechanisms affecting oral absorption of molecules is important in predicting “druggable” molecules for development [4–6]. Prior to intestinal mucosal absorption, the drug molecules undergo harsh conditions in the stomach (high acidity, pH 1.5–3.5) that challenge the stability of the molecules. For example, an ester drug molecule could be subjected to a hydrolysis reaction before entering the small intestine. Therefore, many drug formulations including nanoparticles have been Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00002-8

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designed to protect drug molecules from acidic degradation in the stomach. The formulation of an oral drug can be in solid, solution, or dispersion forms; however, the drug molecules must be soluble in order to cross the absorptive epithelial cells in the small intestine. The drug molecules can cross the absorptive epithelial cells via transcellular or paracellular pathways. In the transcellular pathway, the soluble molecules partition into lumen cell membranes of the intestinal mucosa before entering the intracellular space and finally crossing the basolateral membranes into the bloodstream. In the paracellular pathway, the drug molecules cross the intestinal barrier via the intercellular junctions of the absorptive epithelial cells. Recently, the clinical use of biologic drugs such as proteins (antibodies, enzymes, hormones), peptides, and oligonucleotides has increased. However, these molecules are very difficult to deliver via the oral route due to their physicochemical properties, including size, hydrophilicity, and hydrogen-bonding potential, which prevent them from crossing the intestinal mucosal barrier. Several nanoparticle formulations have been investigated for improving drug delivery of peptides with limited success, and this type of formulation is being continuously improved for future development of peptide oral absorption. Thus, this chapter is focused on describing the organization and structure of the intestinal mucosal barrier as well as the biochemical compositions that are involved in oral drug absorption.

2. Structure of the small intestine The small intestine is approximately 3–5 m in length, with a diameter of about 3 cm; it is segmented into three sections, including the duodenum, jejunum, and ileum (Fig. 1). The duodenum is attached to the lower part of the stomach and receives digestive enzymes from the pancreas. The jejunum, which is the middle segment of the small intestine, has a length of about 2.5 m. Finally, the ileum section is about 3 m in length with structure similar to that of the jejunum and contains various bacteria populations and Peyer’s patches.

2.1 Cellular compositions The lumen of the intestinal mucosal barrier has microvilli structures on their tips (Fig. 1). They are composed of epithelial cells with a column-like structure decorated around them with a few goblet cells. The seams between the absorptive epithelial cells and their interconnections with other cells (e.g., M cells) are connected by “Velcro” or “Zipper” proteins from the cell-cell adhesion family of proteins. The cells in the brush border region, or microvilli, are constantly regenerated by the renewal of embedded proteins. The microvilli provide a large surface area of absorption in which the microvillus is 0.5–1.5 μm in length and 0.1 μm in width. The microvillus membranes

Organization of the intestinal mucosa and barriers to oral drug delivery

Absorpve Epithelial Cell

Absorpve Epithelial Cell

Absorpve Epithelial Cell M Cell

Goblet Cell

Lymphocytes

M Cell

Crypt Blood Vessel Crypt Cell

Lymphac System

Fig. 1 The structure of the small intestine villi, composed of absorptive epithelial, Microfold (M), goblet, and crypt cells. The blood vessels and lymphatic system are found at basolateral side of villi. The M cell is in between epithelial cells where many lymphocytes are crossing the intestinal mucosa. The goblet cells are sporadically located in between the epithelial cells and secret components of the mucus layer. The crypt region at the bottom of villi is populated with crypt cells.

at the lumen have a high protein-to-lipid molar ratio and a high degree of cholesterolto-phospholipid ratio (1.26:1.0 in rat and 1.0:1.0 in mouse). In contrast, the cholesterol-to-phospholipid ratio at the basolateral membrane is less than 0.5:1.0 [7, 8]. The sharp face of microvilli is embedded with a high number of transmembrane proteins, including alkaline phosphatase, ATPase, and glucose transporter. Some proteins are located at the hydrophobic center as well as at the hydrophilic glycocalyx outside the membranes of the microvilli. The crypt region of the lumen is located at the lower region of the villi and is composed of crypt cells that function as secretory cells. Undifferentiated crypt cells can differentiate into villus cells and migrate from the crypt region to the villi. In addition, undifferentiated crypt cells can become goblet, Paneth, and endocrine cells. Structurally, undifferentiated crypt cells are significantly different from differentiated absorptive cells. One clear feature of the crypt region is that it has short microvilli with abundant embedded proteins such as dipeptidase and alkaline phosphatase, with less enzyme expressions compared to regular villi. The enzyme expressions can be used as markers to distinguish absorptive cells and undifferentiated crypt cells. Microfold cells (M cells) are epithelial cells that are found patchily in between absorptive epithelial cells at the lymphoid follicles of the Peyer’s patch, which is populated by

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lymphocytes that can travel across in the intercellular spaces of the epithelial cells. The M cells are connected to the absorptive epithelial cells via tight intercellular junctions. They can transport antigens, particles, and microorganisms from the lumen to cells of the immune system. The M cells have a rectangular-to-oval shape with a microfold structure and are located at the wall of the intestine. Peyer’s patches with M cells are found abundantly in the ileum. At the peak of an M cell, there is a thin cytoplasmic band connecting it to the absorptive epithelial cells. The connecting region has an intercellular cavity between the absorptive epithelial cells. Goblet cells are derived from undifferentiated crypt cells; their nuclei are found in the basal region. The goblet cells have microvilli with a low expression of proteins; however, these microvilli are less uniform than those of absorptive epithelial cells. The mucus layer is produced by the goblet cells, and the majority of goblet cells can be found at the ileum. The components of the secreted mucus are largely glycoproteins with a high content of carbohydrates that are excreted via exocytosis from the cytoplasmic domain of goblet cells. It has been suggested that the function of mucus is to protect the absorptive epithelial cells from bacteria, unwanted particulates, and toxins. It has also been indicated that the mucus can function as another passive diffusion barrier for drug molecules before they encounter the epithelial cells.

2.2 Tight junctions or zonula occludens At the top of the apical side, connecting the epithelial cells is the tight junction or zonula occluden (Fig. 2). The cell membranes of neighboring cells are bridged via cell-cell adhesion molecules such as claudins, occludins, and junction adhesion molecules (JAM). The density of the intercellular membranes prevents molecules from crossing through the tight junctions. Only small ions and molecules with hydrodynamic radius of less than 11 A˚ can permeate the tight intercellular spaces. The tight junction morphology varies from region to region of the intestinal mucosa, depending on cell types along the axis of crypt-villus. The position of the tight junctions in the ileum cells is lower than in the jejunum. In addition, there are more cell adhesion strands found between the absorptive cells at the Ileum compared to those in the jejunum region. The three main cell-cell adhesion proteins at the tight junctions are the claudin family (Cldn), the occludin family (Ocld), and the IgG-like family of JAMs (Fig. 3). In this case, claudins have an important contribution to the tight junction barrier properties [9, 10]. The claudin family is in the tetraspanin family of proteins, with 26 family members in humans [10]. Cldn-1, -2, -3, -4, -5, 7 and -8 are found in the GI tract [9, 10]. These proteins can be divided into “tight or sealing” Cldn and “leaky or pore-forming” Cldn. Cldn-1, -3, -4, -5, -6, -8, -12, -18, and -19 can be categorized as “sealing” Cldn. The “leaky” Cldn includes

Organization of the intestinal mucosa and barriers to oral drug delivery

Apical Side (AP) Paracellular Pathway

Transcellular Pathway

Microvilli

P-gp Efflux Pump

Enzyme Pepde Transporter

CYP Intercellular Juncon

Basolateral Side (BL) Fig. 2 The general structure of absorptive epithelial cells with microvilli or brush borders that are decorated with P-gp efflux pump, peptidase enzymes, and transporters (i.e., PepT1, PepT2). Cytochrome P450 (CYP) metabolic enzymes are also present in the epithelial cells. The transport of drug molecules through the intestinal mucosa from the AP (lumen) to the BL (blood) side can utilize transcellular and paracellular pathways. The paracellular transport of drugs through the intercellular junction is limited by the presence of tight junctions. The transport of drug molecules via transcellular pathway can be modified by P-gp efflux pump. The presence of proteolytic enzymes degrades peptides and protein molecules and CYP can metabolize drug molecules during transcellular transport. Some nutrient molecules such as glucose, amino acid, di- and tri-peptide are carried from the AP-to-BL by transporters.

Cldn-2 and Cldn-15, which have a role in increasing paracellular permeation of sodium ion and water molecules [10]. Cldn proteins have two extracellular loops, extracellular loop 1 (ECL1) and extracellular loop 2 (ECL2), with different functions [11]. ECL1 acts as a paracellular pore control to allow selective permeation of ions and water; the amino acid composition and arrangement contribute to this role [12]. The structural stability of Cldn is influenced by ECL1 via the formation of a disulfide bond by two cysteine residues [11]. The role of ECL2 occurs in cis- and trans-claudin interactions, where the cis-interaction is between two claudins on the same cell, while the trans interaction involves two claudins on opposing cells to generate cell-cell adhesion [13]. The expression and function of tight junction proteins can be influenced by disease states of the GI tract, such as inflammatory bowel disease (IBD), Crohn’s disease (CD), ulcerative colitis (UC) and colorectal cancer [10]. The expression levels of Cldn in the intestines can be influenced by other genes or proteins. As an example, mutant mice with damaged cathepsin have increased levels of Cldn-1 and -2 that cause intestinal neoplasia. The downregulation of β-catenin decreases the Cldn-1 expression in the

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Fig. 3 The structure of intercellular junction between two epithelial cells of the villi containing cell-cell adhesion proteins that connect intercellular junction cell membranes of two opposing cells. The tight junctions are connected by claudins, occludins, and JAMs in the extracellular domain as well as ZO-1, -2, -3, and actin filaments in the cytoplasmic domain. Adherens junctions are decorated with E-cadherins and Nectins that are connected to catenins, ZO-1 and actin filaments in the intracellular space. Finally, desmosomes are comprised of desmocollin and desmoglein and are connected to the cytoplasmic plakoglobin, plakophilin, desmoplakin, and actin filaments.

intestine. In IBD or CD, the expressions of Cldn-1 and -2 in the intestinal epithelium are upregulated, while the expressions of Cldn-3, -5, -8, and -12 are downregulated. In IBD, the expression of Cldn-1 is downregulated at the transmigration region of the epithelium and its localization becomes irregular, causing loosening of intercellular junctions. In colorectal cancer, Cldn-1, -2, -3, -4, -7 are overexpressed while Cldn-8 is suppressed. Thus, these changes could be used as biomarkers for colorectal cancer detection. Pathogenic (e.g., bacteria or viruses) infections of the GI tract influence the function of Cldn on the absorptive epithelial cells. These infections disrupt Cldn interactions and reduce the trans-epithelial resistance (TEER) of the epithelium. Known pathogens that disrupt the tight junction integrity include Escherichia coli (e.g., enteropathogenic E. coli (EPEC), enteroaggregative E. coli (EAEC), enterohemorrhagic E. coli (EHEC)), Yersinia enterocolitica; Salmonella typhimurium; Shigella flexneri serotype 2a; and Giardia lamblia. Some flagellated parasites (e.g., Giardia intestinalis trophozoite) can break down the integrity of the cell-cell adhesion in the tight junctions of the intestinal epithelium. An anaerobic parasitic amoebozoan (e.g., Entamoeba histolytica) produces prostaglandin E-2 (PGE-2) that disrupts the tight junction integrity and increases Na+ permeation. PGE-2 has been shown to cause

Organization of the intestinal mucosa and barriers to oral drug delivery

internalization of Cldn-4 from cell membranes into the cytoplasm. It is interesting to find that some probiotic bacteria (e.g., Bifidobacterium bifidum and Lactobacillus plantarum) could improve the integrity of the intestinal mucosal barrier in a rat model of necrotizing enterocolitis [14]. These probiotic bacteria have been shown to reverse the tight junction dysfunction in Caco-2 cell monolayers. Various molecules have been found to modulate Cldn interactions in the intestinal mucosal barriers, including platelet-activating factor (PAF) and mycotoxins (deoxynivalenol (DON) and ochratoxin A). They downregulate the expressions of Cldn-1 and Zona occludin-1 (ZO-1), resulting in dissociation of claudin-mediated cell-cell adhesion [9, 10]. In contrast, trinitrobenzenesulfonic acid (TNBS) increases epithelial Cldn-1 in actively inflamed intestinal mucosa of mice with ulcerative colitis [10]. However, inflammation-mediated changes of Cldn-1 may vary depending on the nature of the injury, time of onset, and duration of inflammatory response [9]. Moxibustion, fish oil, kaempferol, quercetin, nicotine, butyrate, TGF-α, interleukin 17, and intestinal alkaline phosphatase have been shown to increase the intestinal mucosa integrity and restore Cldn functions in certain pathological conditions. Thus, these compounds could be used to restore dysfunction in the intestinal mucosal barrier [9]. Occludins (Ocld) are in the tetraspanin family and reside in the tight junctions. There are four-helix bundle transmembranes of occludins integrated within the cell membranes [15]. As in Cldn, they have ECL1 and ECL2 loops and four transmembrane helices [15–19]. The N- and C-termini are located in the intracellular space, and they interact with ZO-1, -2, and -3 proteins. As in Cldn, the second loop is important for trans-interaction between Ocld for cell-cell adhesion. JAMs are in the immunoglobulin family and are about 40 kDa in size. JAM protein has a single transmembrane domain with C-terminal intracellular space [20]. JAM-1 and -4 are expressed in intestinal epithelial cells and JAM-4 is also found in renal glomerulus [21]. JAM-1 has a 215 amino acid extracellular domain with two V-type immunoglobin loops; each loop has a disulfide bond [22]. The C-terminal cytoplasmic domain has a 45 amino acid residue and is connected to ZO-1 [23]. JAM-2 is found in lymph and kidney [23]. JAM-3 is expressed at high levels in kidney, brain, and placenta [24]. In the cytoplasmic domain, the C-termini of Cldn, Ocld, and JAMs interact with ZO-1 followed by interactions of ZO-1 with ZO-2 and ZO-3. Cldn connects to ZO-1 at the PDZ1 domain. ZO-1 interacts with Ocld via the guanylate kinase-like (GUK) domain. ZO proteins are in the membrane-associated guanine kinases (MAGUK) family [25, 26]. JAM-1 is linked to ZO-1 via the PDZ3 domain. Phosphorylations of Cldn and Ocld stabilize their localization at the tight junctions; however, dephosphorylation of Cldn and Ocld causes their internalization into the cytoplasmic domain, resulting in the disruption of the tight junctions [27–29].

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2.3 Adherens junctions or zonula adherens The adherens junctions or zonula adherens are directly underneath tight junctions with intercellular spaces 15 to 20 nm wide (Fig. 3). The intercellular space is connected by homophilic interactions of E-cadherins as well as nectins. In the intracellular space, they are connected to the actin filaments. E-cadherin has an extracellular domain with five immunoglobulin repeats called extracellular domain-1 (EC1) to EC5 with the N-terminal of the EC1 domain. The EC5 is the closest to the cell membranes and is connected to a single transmembrane domain followed by the C-terminal intracellular domain. The intracellular domain binds to β-catenin, which is connected to α-catenin and p-120 catenin as well as an actin filaments network of the membrane cytoskeleton [30–32]. In the adherens junctions, E-cadherins interact with each other in a homophilic manner to promote calcium-dependent cell-cell adhesion at the two opposing cells [33–35]. It has been shown that E-cadherin binds to internalin A on Listeria monocytogenes for its transport through the intestinal mucosa. During inflammation, E-cadherin can be utilized by leukocytes via αEβ7-leukocytes for migration through the GI tract [36, 37]. Calcium ions play an important role for E-cadherin interactions, and the presence of calcium cadherin forms a rod-like structure; in contrast, without calcium, E-cadherin forms a globular structure [38]. It has been shown that the cell-cell adhesion property of E-cadherin was diminished in the absence of calcium ions. In a functional cadherin, there are three calcium ions at the interface between two EC domains (e.g., between EC1 and EC2) with DXNDN as a calcium-binding sequence [33, 39, 40]. Two calcium ions bind to two sites with the same Kd values of 330 μM, and the third calcium binds with a Kd of 2.0 mM [33, 39, 40]. Nectins are cell adhesion molecules that belong to the immunoglobulin (Ig) family and are non-calcium-binding proteins that consist of Nectin-1, -2, -3, and -4 [41–43]. The structure of nectin contains three-Ig loops in the extracellular domain followed by a transmembrane domain and a cytoplasmic domain. The first Ig-1 loop is responsible for the trans-dimeric interaction of nectin while the Ig-2 loop is responsible for cis-dimerization to another nectin within a cell membrane [44]. The cytoplasmic domain binds to the PDZ region of the cytoplasmic protein afadin and catenin complex [42, 45].

2.4 Desmosomes or macula adherens Lastly, the desmosomes are approximately 0.2 μm from the basal end of the adherens junctions and contain desmocollins (Dsc) and desmogleins (Dsg) (Fig. 3). Dsc and Dsg belong to the classical cadherin family with five extracellular repeats or domains (EC1–EC5) [46, 47]. They are calcium-binding proteins that can form homophilic and heterophilic interactions. It has been indicated the trans-interaction between Dsc and/or Dsg involves domain swapping using the conserved Trp-2 residue that binds to a hydrophobic pocket of the EC1 on the counterpart protein from the other cell membranes [46, 47]. Their C-terminal

Organization of the intestinal mucosa and barriers to oral drug delivery

domains in the cytoskeleton have additional domains similar to classic cadherins. At their C-termini, Dsg and Dsc have an intracellular anchor (IA) domain and an intracellular cadherin-like sequence (ICS) domain. In addition, Dsg contains an intracellular prolinerich linker (IPL) domain, repeat unit domains (RUD), and a desmoglein terminal domain (DTD) [48]. In the intracellular domain, the C-terminal domain Dsc or Dsg bind to plakoglobin that is connected to desmoplakin. Desmoplakin then binds to intermediate filaments to form the membrane cytoskeletal network. In the epithelial tissue, the intermediate filaments bind to cytokeratin filaments [46, 47, 49].

3. Metabolizing enzymes Enzymes have an important role as metabolic barriers at the brush border and intracellular domain of the absorptive intestinal epithelium (Fig. 2) [50, 51]. The expression of peptidases increases from the upper duodenum to the lower ileum; in contrast, the lowest expression of enzymes is found in the colon. The function of these enzymes is to degrade proteins and carbohydrates to nutrients; conversely, they degrade dangerous molecules (e.g., toxins) to prevent them from crossing the intestinal mucosa into the systemic circulation. Therefore, the metabolic properties of enzymes become an important consideration in the drug absorption process. For example, peptide drugs can be digested prior to absorption, thus limiting the amount of intact peptide that can cross the intestinal mucosal barrier. The variability or genetic polymorphism of the expressed enzymes in the intestinal mucosa affects variability in drug absorption in different individuals. A disease state could also influence enzyme expression and activity. In the ileum, duodenum, and colon some enzymes are expressed on the lumen along with those produced by surrounding bacteria. In the duodenum, many proteases (e.g., trypsin, chymotrypsin, elastase, and carboxypeptidase A and B) rapidly metabolize 30–40% of peptides and proteins that pass through the GI tract [52]. On the brush border, the proteolytic enzymes are normally selective for tripeptides while the cytosol enzymes are selective for dipeptides. The small intestine contains a cytochrome P450 (CYP) superfamily as Phase I metabolizing enzymes (Fig. 2) [53–56]. The sub-family of CYP includes CYP1, CYP2, and CYP3 enzymes, and the enzymes found in the small intestine are CYP1A1, CYP2C, CYP2D6, and CYP3A4. CYP3A4 is the most highly expressed in the intestinal barrier; it is as high as 70% of the available CYP enzymes [57, 58]. Although CYP enzymes have lower populations in the small intestine than in the liver, their activity in the intestine is similar to that of those found in the liver. The highest activity is found at the tips of villus of the middle intestine. Variabilities in enzyme activity are found in different individuals, and this is presumably introduced by foods or drugs that are consumed by the individual [59]. Intestinal CYPs are more responsive to stimulators than are those in the liver. The wide surface area provides easy access for interaction between the enzyme and its substrates, which contributes to the overall enzyme activity. The long residence time of food

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or enzyme substrates in the small intestine offers a high probability of substrate-enzyme interactions for metabolism to occur. In addition, drug trapping in the cytoplasmic domain could contribute to the long residence time of the drug in the intestinal mucosa for metabolism by CYP enzymes. It has been shown that some foods such as grapefruit can inhibit the activity of CYP3A; thus, some drugs cannot be taken with grapefruit because it can change the drug absorption properties [59]. Phase II enzymes in the intestinal barrier are metabolizing as well as conjugating enzymes [53, 60, 61]. Conjugating enzymes covalently connect a drug molecule with another moiety to form a drug conjugate that becomes a substrate for the efflux pumps found in the lumen. High expressions of Phase II enzymes (e.g., glutathione-S-transferase, N-acetyltransferase) are found in the intestinal mucosa epithelium [60, 61].

4. Efflux pumps The absorptive cells have efflux pump transporters (e.g., P-glycoprotein, P-gp) that can prevent molecules from crossing the cell membranes of the intestinal mucosa even if the drug molecules effectively partition into the cell membranes (Fig. 2) [54, 62, 63]. The transporter activities are similar to those of the multidrug resistance protein that prevents anticancer drugs (i.e., doxorubicin and paclitaxel) from entering the cancer cells [64]. P-gp proteins are found on the apical side of the intestinal mucosa layers that polarize the absorptive epithelial cells. P-gp is in the family of ATP-dependent multidrug resistant (MDR) transporters along with multidrug resistance-associated proteins (MRPs), and breast cancer-resistant protein (BCRP) [65]. MDR1 and MDR2 are two members of the P-gp family; however, only MDR1 has drug transport activity. The size of human P-gp is between 140 and 190 kDa, depending on glycosylation levels. P-gp transporters are embedded in the cell membranes of the absorptive epithelial cell using six transmembrane domains. The levels of P-gp expression and polymorphism can vary among individuals depending on age and gender, and these variations could contribute to drug absorption variability among individuals. P-gp transporters have broad substrate specificity with commonly known substrates, such as anthracyclines, paclitaxel, rhodamine 123, topotecan, and HIV protease. It has been suggested that these pumps recognize hydrophobic molecules with log P octanol/ water > 1.0 with substrate characteristics to be aromatic or planar neutral structure (e.g. Phenytoin) as well as positively charged groups such as tertiary amines [66–68]. However, negatively charged molecules such as methotrexate have been shown to be P-gp substrates under certain conditions. In Caco-2 cell monolayers, basolateral-toapical transport of fexofenadine (FXD) is five times higher than apical-to-basolateral, and the transport of FXD can be altered by verapamil, suggesting that FXD transport involves an P-gp efflux pump [69].

Organization of the intestinal mucosa and barriers to oral drug delivery

Inhibition of P-gp enhances the intestinal drug absorption, and certain types of foods influence drug absorption due to inhibition of P-gp. Verapamil, valinomycin, cyclosporin A, quinidine, LY335979, and other molecules are inhibitors of P-gp. Pharmaceutical excipients such as polysorbate 80, cremophore, polyethylene glycols, and pluronic acid also increase oral drug absorption because they inhibit P-gp activity [70–72]. Cyclosporine D (SDZ PCS833) and dexverapamil as inhibitors of P-gp have been investigated as adjuvants to improve oral bioavailability of drugs; the hope is that these molecules inhibit P-gp without retaining their original target receptor bioactivities such as cyclosporin A and verapamil [66, 73]. The mechanisms of action of P-gp transporters to expel their substrates have been investigated using its X-ray structure. The first proposed mechanism is called the “hydrophobic vacuum cleaner” mechanism; in this mechanism, the hydrophobic substrates are embedded in the plasma membranes followed by binding to P-gp; the substrates are then pumped out of the cell membranes into the extracellular space. The second mechanism is the “Flippase” mechanism, in which the substrate binds to P-gp at the inner cell membranes, followed a flipping motion of the P-gp to move the substrate into the outer cell membranes where the substrates diffuse into the extracellular space. Finally, the “pore” mechanism was proposed in which P-gp transporters create pores in the cell membranes to allow diffusion of drug molecules from the membranes into the extracellular space. The broad substrate structure recognition can be explained by high concentrations or high populations of drug molecules in the cell membranes that can bind to P-gp as substrates; thus, even if the substrate has low affinity to P-gp, a fraction of the substrate still binds to P-gp for efflux from the cell membranes into the extracellular space.

5. Transporters for transcytosis Molecules can be shuttled from the gut lumen into the bloodstream by transporters. The morphology and biochemical composition of lumen is different from that of the basolateral membranes while both contain transporters. Some of these are di- and tri-peptide transporters where the activity of the transporter is energy-dependent and often pH-dependent. The important characteristic of the transporter is that it can be saturated by its substrates. Peptide transporters such as human intestinal peptide transporter 1 (hPepT1) are found on the tip of the villus. hPepT1 is a proton-dependent transporter whose population increases from duodenum to ileum. hPepT1 transports oligopeptides, and has been used previously to improve absorption of peptide-based prodrugs [74, 75]. Ibuprofen, a non-steroidal antiinflammatory drug, has been shown to be a non-competitive inhibitor of hPepT1 in the Caco-2 cell line. In the basolateral side, PepT2 transporter carries substrates from the intracellular to the basolateral side (i.e., blood side) of the absorptive epithelial cells. Other basolateral transporters include sodium-dependent amino acid, GLUT2 hexose, and folic acid transporters.

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Some hydrophilic molecules such as folates and methotrexate are orally absorbed because they are carried by active transporters such as reduced folate carrier (RFC) and membrane folate binding protein (mFBP). As nutrients, amino acids in the lumen cross the intestinal mucosa into the bloodstream via amino acid transporters. Peptide transporters (e.g., hPepT1) have been used to improve oral absorption of valacyclovir and other molecules by conjugating the drug to the valine amino acid [76, 77]. Similarly, dipeptide transporters can carry dipeptide-like drugs such as beta-lactam, captopril, enalapril, lisinopril, renin inhibitor, and carnosine across the intestinal barrier. The organic cations are shuttled by and OCT1 and OCTN2 transporters while the organic anions are carried by OATP1A1 and OATP1B3 transporters across the intestinal mucosa. Finally, glucose is absorbed by the gut into the systemic circulation by Glut1 transporters [78].

6. Mechanism of crossing the intestinal mucosal barrier 6.1 Passive diffusion through transcellular pathways Most hydrophobic molecules cross the intestinal mucosal barrier by passive diffusion via the transcellular pathway. In this case, the drug molecules have to effectively partition into the cell membranes of the epithelial intestinal mucosal barrier (Fig. 2) [1, 2]. Thus, the drug has to have the appropriate physicochemical properties with sufficient solubility [5, 6]. One potential guide to effective transcellular passive diffusion is the Lipinski’s rule of five [79, 80]. Although this rule has its own weaknesses, it is a good way to predict whether a molecule can cross the cell membranes via passive diffusion. In this case, the rule stated that for a molecule to passively diffuse through cell membranes, it must have (a) octanol: water partition or Log P less than 5.0, (b) molecular weight less than 500 Daltons, (c) fewer than 5 hydrogen bond donors, and (d) less than 10 hydrogen bond acceptors [79, 80]. Although a drug molecule can effectively partition to the cell membranes, it can also be inhibited from crossing the intestinal mucosal barrier by P-gp efflux pumps. In addition, drugs containing an amino group(s) can undergo intracellular sequestration in the endosomes and lysosomes; this is due to the acidic pH nature of the cellular compartments [81, 82]. At acid pH, the amino group is protonated to make a positively charged drug inside the endosome; the charged drug cannot freely partition into endosome membranes. Therefore, the molecule is sequestered in the compartments of absorptive epithelial cells and is prevented from crossing into the systemic circulation. As examples, propranolol and amodiaquine have been shown to be trapped inside membrane compartments (e.g., endosomes, lysosomes) of Caco-2 cell monolayers [83]. Similarly, while methotrexate is transported by RFC and mFBP, it can be trapped inside the absorptive epithelial cells (e.g., Caco-2) because the MTX undergoes polyglutamation with a high number of negative charges [84, 85].

Organization of the intestinal mucosa and barriers to oral drug delivery

6.2 Passive diffusion through paracellular pathways Passive diffusion through paracellular pathways is limited to ions and small hydrophilic molecules such as mannitol (Fig. 2) [1, 2]. The presence of tight junctions prevents large hydrophilic molecules such as peptides and proteins from passing through the paracellular pathways of the intestinal mucosa. Because of their physicochemical properties, peptides and proteins cannot readily cross via transcellular pathways; they also cannot pass through paracellular pathways because of the limited porosity of the tight junctions. One way to permit large molecules to cross the tight junctions is by temporarily disrupting the tight junctions to increase their porosity. Several adjuvants have been investigated in formulations to improve oral absorption of drug molecules via paracellular pathways. Tight junction disruptors such as cationic polymers (chitosans, heparins, and gelatins) [86–89], chelating agents (ethylenediaminetetraacetic acid) [86, 90, 91], plant-derived materials (herbal and Aloe vera gel) [86, 92–95] and toxins (zonula occludens toxin) [86, 96, 97] have been shown to improve paracellular permeation of molecules. Unfortunately, some of these have some toxicity to the absorptive cells. AT-1002 peptide was successfully used to enhance the paracellular delivery of low molecular weight heparin in Caco-2 cell monolayers [1]. A large protein, albumin (66.5 kDa), can cross the Caco-2 cell monolayers with the help of PN-78 and PN-159 peptides as tight junction modulators [98]. Several peptides (e.g., HAV and ADT peptide) derived from the EC1 domain can modulate cadherin-cadherin interactions in the adherens junction of the Caco-2 and MDCK cell monolayers. This peptide modulation enhances paracellular permeation of small hydrophilic molecules such as 14C-mannitol. Both peptides can lower the TEER values of MDCK monolayers, suggesting that they can modulate the intercellular junctions. Chitosan has been used as an intestinal absorption enhancer via paracellular pathways. It has been suggested that the positive charges in chitosan have a role in chitosan interaction with the epithelial cells of the intestinal mucosa to increase the paracellular porosity. Its interaction with cell membranes induces reorganization of the tight junction membranes to enhance drug paracellular penetration [99]. Using the same mechanism, N-trimethyl chitosan chloride (TMC) has been shown to improve absorption of large hydrophilic compounds across Caco-2 cell monolayers [87, 88].

7. Models to study intestinal absorption 7.1 In vitro models Several in vitro cell culture models have been developed for rapid evaluation of the intestinal absorption properties of potential drug molecules. The Caco-2 cell monolayer is one of the most well-studied models for evaluating oral absorption, and this in vitro model can reasonably predict oral drug absorption with good correlation to the results obtained

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using the in vivo model [100, 101]. Caco-2 cells are derived immortalized human epithelial colorectal adenocarcinoma cells and they have characteristics comparable to absorptive epithelial cells of the small intestine [100–102]. This similarity includes the polarized nature of Caco-2 cell monolayers with apical (AP) side resemblance to the brush border as well as the basolateral (BL) side with all its components [1, 101]. This model contains active transporters, efflux pumps, and appropriate metabolic enzymes mimicking the in vivo conditions [51, 103].

7.2 In vivo models The most logical progression after in vitro studies is to evaluate intestinal drug absorption using in vivo models in small rodents such as mouse and rat. The use of animal models provides ways to investigate the formulations, excipients, and adjuvants that are effective in delivering potential drugs in the GI tract. Besides rodents, other animals, including rabbits, dogs, sheep, monkeys, and non-human primates have been used to study oral drug absorption [104]. The use of an in vivo model is to confirm results from in vitro studies as an intermediate progression to clinical trials [104, 105]. Although there is an established in-situ rat intestinal perfusion method, a non-invasive in vivo method is more favorable to mimic the situation as close as possible to the human oral absorption [104, 106]. The advantage of in vivo studies over in vitro is that, after oral absorption, the drug pharmacokinetics and biodistribution can be followed. In in vivo studies, the drug can be monitored not only in the stomach, small intestine, and colon, but also in the blood, lymphatic system, and other organs [106, 107]. In addition, the side effects or organ toxicity of the drug as well as the immunogenic response can be monitored in the in vivo system [106, 108]. However, oral delivery of peptides and proteins still remains challenging for various reasons, including unfavorable physicochemical properties and low stability.

8. Conclusions Because oral delivery is one of the most convenient methods, many drug candidates are investigated for transport via this route. However, the physical, biochemical, and biological nature of the intestinal mucosa does not allow all molecules to cross this barrier. Thus, understanding various aspects of the intestinal mucosal barrier is essential in designing drug formulations that can be successful in delivering the desired drug orally. Although many peptides, proteins, and other biologic drugs are challenging to deliver orally, many methods including development of polymer and nanoparticle delivery systems have been investigated for improving their intestinal transport. Several methods have also been investigated to improve intestinal mucosa delivery of peptides and proteins through the paracellular pathways; so far, these methods have had limited success in vivo. Although they have met with various challenges, scientists from different fields are working diligently to

Organization of the intestinal mucosa and barriers to oral drug delivery

develop novel methods to overcome the difficulties in delivering drugs across the intestinal mucosa for oral drug delivery.

Acknowledgments We acknowledged funding for our research from P30-AG035982 from the KU Alzheimer’s Disease CenterNational Institute on Aging (NIA) and R01-NS075374 from the National Institute of Neurological Disorders and Stroke (NINDS), National Institutes of Health. We would like to thank Nancy Harmony editing the manuscript for this Chapter.

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CHAPTER 3

Nanomaterials for oral drug administration Serena Bertoni, Nadia Passerini, Beatrice Albertini

Department of Pharmacy and BioTechnology, University of Bologna, Bologna, Italy

1. Introduction Conventional formulations including both liquid (solutions and suspensions) and solid (granules, capsules and tablets) dosage forms have represented a benchmark for oral drug products over the years. Since the ‘90s, the pharmaceutical industry has recognized nanotechnology as one of the most powerful tool to overcome the issue of poorly watersoluble drugs by reducing drug particles to nanometer size [1]. The terms “nanoparticles” or “drug nanocrystals” have been used to describe nano-sized particles consisting entirely (or for the most part) of the active pharmaceutic ingredient (API). Due to this approach, it is possible to overcome issues related with low solubility and/or dissolution rate by drastically enhancing the surface area of drug particles through nanosizing. Starting from Gris-PEG, the first oral product to be approved by the Food and Drug Administration (FDA) containing the active compound in nano-sized particles, many other oral products based on nanocrystals have been launched [2]. In the wake of this success, the combination of material engineering with the knowledge in biology and pathophysiology has led to the increasing development of therapeutics based on nano-sized materials, or nanomedicines [3]. The research has continued mainly focusing on nanomedicines for parenteral delivery, with considerable efforts dedicated to tumor targeting, chasing Paul Ehrlich´s idea of a “magic bullet” that transports the drug directly to the targeted site, bypassing healthy tissue [4]. On the other hand, more and more colloidal drug carriers designed for other routes of administration, including the oral way, started to be investigated. To date, a wide variety of oral nano-scaled drug delivery systems have been developed. The final goal is the improvement of the therapeutic outcome of already marketed oral drugs with bioavailability issues, and the offering of novel efficient oral formulations for newly discovered compounds or for APIs already marketed by different routes. The number of published studies on nanoformulations for oral drug administration has increased exponentially over the past decade, as illustrated in Fig. 1. To date, these systems include nanocrystals, inorganic nanocarriers, polymeric-based nanocarriers

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© 2020 Elsevier Inc. All rights reserved.

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Fig. 1 The increasing trend of research about nanomedicines for oral drug delivery belonging to the different categories (nanocrystals and nanosuspensions, polymeric-based nanocarriers, lipid-based nanocarriers and inorganic nanocarriers), as reflected by increasing publications in Scopus from 2000 to 2018.

and lipid-based nanocarriers. The majority of the developed nanoformulations belongs to the last two categories, and specifically more than 25% of publications refer to formulations based on polymeric materials, while lipid-based nanocarriers are the protagonist of more than 50% of the studies. In fact, nanocarriers are versatile delivery systems, as their physiochemical properties can be modulated by modifying their composition, size, shape and surface properties (e.g., surface charge, coating or attachment of targeting moieties) [5]. The API can be attached, absorbed, physically entrapped or dissolved into the nanosized vehicle. Differently from the conventional oral dosage forms, nanocarriers offer the possibility to customize the nanoparticle surface with ligands to target a particular region of the gastrointestinal tract (GIT) or to provide a special feature to the nanocarrier [6]. Furthermore, the exploitation of endogenous triggers (e.g., a particular pH) related to the diseased site has led to the development of nanocarriers based on stimuli-responsive materials [7]. Overall, nanoformulations find applications in oral drug delivery for various reasons and can play different roles: (i) Vehicles for the encapsulation and protection of molecules susceptible to degradation or instability in the GIT, as in case of biomacromolecules, peptides, nucleic acids. (ii) Means to enhance the apparent aqueous solubility of poorly soluble compounds and enable the formulation of hydrophobic drugs. (iii) Carriers that can penetrate through intestine epithelium and release the drug in the blood stream or increase drug absorption by other mechanisms (e.g., inhibiting the efflux pumps or increased mucosal contact). (iv) Carriers able to target particular delivery sites in the GIT.

Nanomaterials for oral drug administration

2. The rational design of nanocarriers for oral drug formulation The physiology of the GIT is complex and presents a wide range of environments characterized by different fluid volumes, composition and pH, the presence of food and digestive enzymes, all factors that can affect drug stability. Moreover, the intestinal epithelium covered by mucus layer acts as a tight barrier hindering drug absorption. Finally, other obstacles for oral bioavailability are represented by first-pass metabolism and susceptibility to efflux mechanisms. All these barriers constitute a challenge for the achievement of a reliable and efficient delivery of drugs via oral route. This is particularly true for some APIs, which present unfavorable physical-chemical properties when administered orally, as for example biologic drugs. The oral delivery of macromolecules such as proteins, peptides and nucleic acids remains a major challenge due to their poor stability in the GI environment, as well as the extremely low absorption caused by their high molecular size and hydrophilicity. These therapeutics are currently restricted to parenteral administration, which impose considerable limitations from the patient and the industrial point of views. Completely different are the issues of small molecule drugs with poor water solubility, belonging to the biopharmaceutical classification system (BCS) class II (poorly soluble, permeable) and class IV (poorly soluble and poorly permeable). For these compounds the low aqueous solubility and intrinsic dissolution rate are the major factors leading to poor bioavailability [8]. It is well known that, in recent years, the fraction of “practically insoluble” molecules constitutes about 70% of the compounds in development [9]. Finally, the formulation of medicines for diseases localized in specific regions of the GIT, such as the colon, represents another key task in the field of oral drug delivery systems. The achievement of an effective local drug delivery in the GIT is complicated due to the distal location of region itself and to difficulties in avoiding the systemic absorption, which causes the adverse effects in long-term treatments. Since the development of nanotechnology, the pharmaceutical research has seen in nanomedicine the potential to address and overcome these challenges. However, the development of oral nanoformulations is more complex than traditional ones, and the way from the formulation to the clinic is long and challenging. Although most drug products manufactured worldwide are intended for oral application, nearly all of the FDAapproved nanomedicines are administered by intravenous injection. A good starting point in the direction of successful oral nanomedicines consists on the development of efficient, robust and feasible formulations. To this aim, multiple factors should be considered and the formulators should have knowledge about: (i) drug loading and release, (ii) biocompatibility and biodistribution, (iii) physical and chemical stability of the formulation, (iv) efficacy, and (v) toxicity. First of all, not all drugs are suitable for being formulated in nanosystems, especially the less potent drugs that require high doses, because typically nanocarriers could not achieve very high drug loading values. Notwithstanding that, a good drug-loading capacity along with high encapsulation efficiency are

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still important prerequisites for a successful nanoformulation. Moreover, stable drug loading values should be maintained during the lifetime of the product (from production until administration to the patient). The nanoformulation must be stable during manufacturing and during storage, and attention should be placed to evaluate the long-term “pharmaceutical” stability of the formulation. Factors affecting the long-term stability of the nanocarrier include reduction of the amount of therapeutic molecule due to leakage or chemical degradation and changes in the nanocarrier structure. In case of liquid formulations (e.g., nanosuspensions), common problems are aggregation, sedimentation and agglomeration of nanoparticles, whereas solid components might undergo changes in crystalline state. Strategies to address different stability issues should be tailored according to the specific drug, the therapeutic requirements and the final dosage form [10]. Upon oral administration, the fate of the nanocarrier is determined by its interaction with the physiological environments that it will encounter during treatment. The control over the nanocarrier’s physico-chemical properties is fundamental, as these properties strongly influence the interplay between the nanomaterial and the GIT. The formulation should retain suitable stability in the GIT conditions and the main factors to consider are chemical stress (i.e., pH, hydrolytic enzymes), residence time and GI fluids volume [11]. After dispersion into the GI fluids, the nanocarriers might be digested (e.g., lipid-based formulations), degraded by other mechanisms (e.g., biodegradable polymeric carriers) or maintain their intact structure (e.g., metallic nanoparticles). In the intestinal tract, the nanoparticles might be immobilized at the mucus layer (mucoadhesion) and eventually come in contact with the intestinal cells. Here, the nanocarrier is supposed to release the drug, which can act locally (e.g., corticosteroids or antibiotics for inflammatory intestinal diseases) or systemically, which implies the absorption through the intestinal membrane. In general, the two main mechanisms of drug absorption are paracellular (passage between the cells) and transcellular (passage through the cells) transports. Transcellular passive diffusion is generally limited to lipophilic molecules with molecular weight below 700 Da, and hydrophilic and/or bigger molecules cannot pass through this pathway. On the other hand, paracellular route allows the transport only of small hydrophilic compounds (molecular weight 220

Spherical

Anionic

NA

[157]

BSA, bovine serum albumin; CMC, carboxymethylcellulose; CS, chitosan; CsA, cyclosporine A; DSNPs, dendritic silica nanoparticles; TPGS, d-α-tocopherol polyethylene glycol succinate; HyA, hyaluronic acid; HPMC, hydroxypropyl methylcellulose; LbL, layer-by-layer; MSNs, mesoporous silica nanoparticles; NP, nanoparticle; NLCs, nanostructured lipid carriers; NA, not available/not applicable; PLGA, poly(lactic-co-glycolic acid); PLA, poly(lactide); PAA, polyacrylic acid; PEG, polyethylene glycol; PL, polylysine; PS, polystyrene; SNEDDS, selfnanoemulsifying drug delivery system; SLNs, solid lipid nanoparticles; SPION, superparamagnetic iron oxide nanoparticles; TMC, trimethylchitosan.

Size, shape and surface charge considerations

diameter microparticles, respectively [160]. Interestingly, in another study, small silica anionic nanoparticles (99%), depending on modular biomolecule templating. Bifunctional amphiphilic peptides designs (e.g., SurSi peptides (AcMKQLAHSVSRLEHA RKKRKKRKKRKKGGGY-CONH2)) are implemented mainly for stabilizing hydrophobic drugs such as curcumin NPs. They induce biosilicification at the surface of the nanodrug particle, forming drug-core silica-shell nanocomposites. This technique is highly adaptable for various hydrophobic cargos. Another example of amphiphilic peptide design includes the CAMSi peptide (CH3(CH2)7COMKQLADSLHQLARQVSRLEHA RKKRKKRKKRKKGGGY-CONH2), which consist of three modules (surface-active module (MKQLADS LHQLARQVSRLEHA), Si module (RKKRKKRKKRKKGGGY-CONH2)), and a hydrophobic tail module (CH3(CH2)7COd). This CAMSi peptide is used to stabilize hydrophobic paclitaxel NPs, which showed to be superior as compared to SurSi peptide. The drug loading efficiency of these drug-core-silica-shell nanocomposites were calculated by thermogravimetric analysis profiles of silica NPs induced by CAMSi (Si NP), paclitaxel-core-silicashell nanocomposites (PTX NC), curcumin-core-silica-shell nanocomposites (CCM NC), paclitaxel (PTX), and curcumin (CCM). The drug loadings were calculated from the weight loss during heating [239]. Amino-modified chiral mesoporous silica nanoparticles (Amino-CMSN) were documented to improve the solubility of indomethacin [240]. The loading was assisted via hydrogen bonding, which resulted in the conversion of indomethacin from crystalline to amorphous form, thus improving the dissolution of the drug [240]. Similarly, in another study, carboxyl group functionalized mesoporous silica nanoparticles loaded with doxorubicin showed higher dispersity and stability in an aqueous solution [241].

Size, shape and surface charge considerations

In addition, these particles exhibited pH-sensitivity, with accelerated release of doxorubicin and higher cellular internalization observed in acidic conditions [241]. The drug-carrier complex formation in nano-based formulations is extremely important, and depends on the physicochemical properties of both the drug and the carrier. Ideally, molecules with straight or branched chain hydrocarbons, aldehydes, ketones, alcohols, organic acids, fatty acids and polar compounds, such as halogens, oxyacids, and amines, promote the process of complexation [242]. The dispersion of drug in the carrier mainly depends on the solubility and type of interaction between them. Similarly, the release amount of dispersed drug will depend on the hydrophilic/hydrophobic nature of the carrier. Hydrophilic carriers usually release the drug at higher rates than hydrophobic carriers [243]. For instance, piperidine was loaded into solid dispersions of hydrophilic carriers made of sorbitol, polyethylene glycol and polyvinyl pyrrolidone K30. This study showed greatly improved piperdine dissolution properties when compared to the native drug, with releases of 70%, 76% and 89% of the drug from the carrier, respectively [244]. Hence, a detailed understanding of the dynamics of the nanocomplexes is necessary to develop formulations with active and controlled release of the therapeutic payloads. Another attractive inorganic material for drug delivery is the porous silicon (PSi), which is produced by an electrochemical synthesis. It allows the construction of tailored pore sizes and controllable volumes in the scale of microns to nanometers. The surface of PSi NPs can be easily modified, allowing for the regulation of the drug loading and drug release rate [245].

4. Summary In spite of its tremendous potential, the oral delivery of nDDS is associated with unaddressed challenges that are essential for the development of a clinical product. Research based on nanoparticle drug carriers could facilitate the development of innovative platform technologies for an effective transport and controlled release of drugs in various disease conditions. Nanoencapsulation offers numerous advantages such as, for example, improvement in drug solubility, distribution, release, and bioavailability. Nanocarriers can be tuned towards favorable physiochemical properties in terms of size, shape (rod, sphere, branched or layered) and surface properties (charge, functional groups or targeting moieties). Many of such nanocarriers have been either approved or present in various phases of clinical trials. Thus, continuous improvements in the development of nDDS can have substantial implications in the cellular uptake and bioavailability of the therapeutic cargo. The translation of nDDS to the preclinical and clinical phases, and commercialization requires research development and characterization to ensure the quality, safety and efficacy of the final products. Size, shape and charge variations of the nDDS should, therefore, be primarily considered. Moreover, drug biodistribution and drug release are also

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important when evaluating the stability of drug nanocarriers in different storage conditions. The size stability is especially critical for micro- and nano-scale drug delivery systems. In this sense, novel or modified nanoparticle characterization techniques should be applied for the development of the next generation of nDDS. In vitro models alone lack the ability to provide quantitative data with respect to drug absorption, distribution, metabolism and excretion. Therefore, animal models should be considered based on the physiological and biochemical similarities with humans. Hence, the appropriate selection of experimental models along with their applications is essential to understand the pharmacokinetics and pharmacodynamics of drugs.

Acknowledgments Authors would like to acknowledge the National Health and Medical Research Council for funding Fellowships to AP and TK. Authors also acknowledge funding received from the School of Pharmacy, The University of Queensland.

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CHAPTER 6

Modified drug release: Current strategies and novel technologies for oral drug delivery Sarah J. Trenfield, Abdul W. Basit

UCL School of Pharmacy, University College London, London, United Kingdom

1. Part 1: Evolution of oral modified release dosage forms Solid oral dosage forms (SODFs) have formed the mainstay of therapeutic practice for over 4000 years, with the earliest reports of pills dating back to 1500 BCE. Crude pills were traditionally composed of medicinal substances such as saffron, myrrh, cinnamon and willow bark, which were incorporated into viscous, sticky materials (such as honey, dough or grease) and spheronised to create small, rounded pills. The Greeks termed these early SODFs “katapotia,” translating directly to “something to be swallowed,” but it was the Roman scholar Pliny (23–79 AD) who coined the more modern term: pilula (pill). During the medieval ages, Rhazes (commonly referred to as the “Father of Pill Coating”) introduced the concept of coating pills with gum mucilage to prevent pills from sticking together and to overcome the offending taste of the medicine. Avicenna (980–1037 AD) is credited with introducing coating of pills with silver and gold into pharmacy practice and, in 1866, a Philadelphia pharmacist was the first to manufacture pills with a sugar coating. The fact that certain coatings were slower to dissolve than others was first noted by Proctor, who believed and stated the following: The reason why pills occasionally pass through the stomach undissolved must be sought for in the state of the patient or in the composition of the pill, rather than in the nature of the coating…

Due to this belief, the use of gastric insolubility as a basis for delayed release was not apparent for 20 more years to come, until Dr. Paul Unna in 1884 introduced keratin-coated pills that remained intact in the stomach and disintegrated in the small intestine, founding the very first delayed-release dosage forms. These efforts represented the very first steps to regulate the release of drug within the gastrointestinal (GI) tract. In the 1930s, the use of wax, vegetable gum, or other enteric material coatings as a method to regulate hydration and drug release were first introduced. However, it was not until after the 1950s that the emergence of more sophisticated extended or timed release oral drug products were produced in North America, enabling for a more precise control over the rate and location of Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00006-5

© 2020 Elsevier Inc. All rights reserved.

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drug delivery. Such developments established a new era of modified release dosage forms, extending through to modern day practices. Traditional immediate release (IR) drug products still account for the mainstay of pharmaceutical preparations available on the market (prescription and over-the-counter). However, a major challenge with the majority of the conventional IR preparations is that they require administration multiple times per day to achieve and maintain the drug levels within the therapeutic range. In turn, this causes a somewhat undesirable “seesaw” effect of drug level fluctuations in the body (Fig. 1). Critically, the advent of modified release drug delivery systems provide a host of benefits compared with conventional immediate release preparations, including: • Reduced frequency of dosing: Plasma concentration can be maintained over an extended period of time, extremely valuable for improving medication adherence for patients with chronic conditions. • Reduction or avoidance of side effects due to high plasma drug concentrations or “dose dumping.” • Improved control of therapeutic drug concentration, particularly important for patients that suffer from breakthrough symptoms. • More cost effective manufacturing due to a reduced number of dosage forms required per patient compared with its immediate release counterpart. To date, modified release dosage forms refer to three main types of oral drug product release pathway: 1. Extended-release: In general, these oral dosage forms should exhibit at least a twofold reduction in dosage frequency as compared to the conventional immediate-release drug product. Examples of extended-release dosage forms include sustained-release, controlled-release, and long-acting drug products.

Fig. 1 Systemic drug levels following oral administration of (A) a conventional immediate release drug product; (B) a zero-order controlled release profile and; (C) a first order sustained release profile.

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2. Delayed-release: A dosage form that releases a discrete portion or portions of drug at a specific time following administration. Enteric-coated dosage forms are common delayed-release products (e.g., enteric-coated aspirin and other steroidal products). 3. Targeted-release: A dosage form that releases drug at or near the intended physiologic site of action or drug absorption. Targeted-release dosage forms may have either immediate- or extended-release characteristics. Examples include gastroretentive devices and colonic drug delivery systems. The following information in this chapter will seek to provide an overview on the common strategies used for extended release, delayed release and targeted release formulations, and provide an update on the latest, cutting edge technologies that are being adopted in oral modified release dosage forms.

2. Part 2: Current oral formulation strategies for modified release 2.1 Extended release formulations Despite the multitude of marketed oral extended release pharmaceutical products, there are actually only four main mechanisms for extending drug release: dissolutioncontrolled, diffusion-controlled, osmotically-controlled and ion exchange methods. To date, the majority of controlled release formulations are designed based on one or combination of these mechanisms. 2.1.1 Dissolution-controlled formulations Dissolution-mediated controlled release formulations can be broken down into two main types: reservoir and matrix systems (Fig. 2). Matrix dissolution systems are the most commonly used technique for controlling release, and involve the API being homogeneously distributed throughout a polymer matrix. As the polymer matrix dissolves (typically via an erosion-mediated process), drug molecules are released into the external environment [1]. In this particular system, the size of the formulation (and hence matrix) decreases over time, thereby resulting in a non-linear drug release. Reservoir systems involve a drug core that is coated with a rate-limiting polymer. Using this system, drug release is determined by the thickness and the dissolution rate of the polymer membrane surrounding the drug core. Once the coated polymer membrane dissolves, the drug will release similarly to an immediate release preparation. It is common for small multiparticulates to be based on this approach, which can be either compressed into tablets (multi-unit particulate systems; MUPS) or filled into capsules, particularly used for taste-masking applications in pediatrics. However, it is worthy to note that the benefits of this system only remain as long as the coating integrity is maintained during the manufacturing and handling processes. This is particularly challenging for MUPS systems which require the application of compression force, in turn potentially affecting the morphology of beads and rupturing brittle coatings (such as ethylcellulose);

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Fig. 2 The two main dissolution-controlled release drug delivery systems; reservoir and matrix systems.

Fig. 3. Due to this risk, this approach is not favorable for tablet formulations due to the potential for dose dumping in vivo. By adjusting the polymer coating thickness, or selecting a polymer that dissolves only at a specified pH, it is possible to design a specific and targeted release profile. 2.1.2 Diffusion-controlled formulations Diffusion-controlled formulations involve the API diffusing through a polymer membrane (reservoir systems) or matrix (monolithic systems) to be released (Fig. 4) [2, 3]. The reservoir systems can further be classified into non-porous and micro-porous systems; the non-porous systems involve API molecules diffusing through an insoluble polymer membrane into the external aqueous environment. Conversely, micro-porous reservoir systems involve the API molecules being released via diffusion through micropore channels, formed upon contact of the polymer coating (combined with poreformers) with water. Diffusion-controlled monolithic systems can also be further classified based on the concentration of loaded drug, termed either monolithic solutions or monolithic dispersions. A monolithic solution is created via API loading by soaking a polymer matrix in a

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Fig. 3 Potential effects of compaction pressure on coated multiparticulate systems.

Fig. 4 Diffusion-controlled release drug delivery systems comprised of insoluble polymers: reservoir and monolithic systems.

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drug solution, in which the drug concentration inside the matrix is not higher than the drug solubility. If the drug loading is higher than the drug solubility, the system is called a monolithic dispersion. Hydrophilic matrix systems are commonly used for diffusioncontrolled release systems (commonly termed swellable soluble matrices). This system comprises an API dispersed within a water-swellable hydrophilic polymer. Upon exposure to an aqueous environment post-administration, the polymer material begins to swell producing a gel matrix. The gel then allows drug release via dissolution or erosion of the gel matrix, enabling drug release into the external fluid. Conversely, insoluble polymer matrix systems can be used whereby the drug is dispersed within an inert polymer. The structure has previously been likened to that of a sponge, whereby water can enter via micro-channels into the matrix, subsequently dissolving the drug and enabling release. The rate of drug release is controlled by three main factors: (1) the tortuosity of the microchannels; (2) the pore size; and (3) the number of pores. 2.1.3 Osmosis-based formulations Osmosis involves the spontaneous and natural movement of water through a semipermeable membrane into a solution of higher solute concentrations. The elementary osmotic pump was first developed in the 1970s designed to enable a zero-order drug release over prolonged periods of time. Favorably, these dosage forms are able to exhibit an extended release profile, independent of gastric pH or hydrodynamic conditions [3, 4]. The oral dosage form involves a drug-loaded osmotic core, which is surrounded by a semi-permeable membrane with a laser hole drilled into the membrane surface (Fig. 5A). Following administration, water penetrates through the semi-permeable membrane into the drug core, dissolving the drug particles and causing drug release via the orifice in the membrane, with the resulting drug solution being delivered at a proportional rate to water entering the tablet. One of the main challenges with the elementary osmotic pump system is that it is only suitable for the delivery of water-soluble drugs. As such, this spurred on the development of an osmosis-based formulation that could be used to control release of poorly watersoluble drugs, known as the “push-pull” osmotic systems (Fig. 5B). The push-pull system is comprised of a bilayer core; one section contains an osmotically-active drug reservoir and the other is comprised of a hydrophilic expanding compartment. Following administration, water migrates across the semipermeable membrane causing twofold effects: (1) creation of a drug suspension or solution and (2) a swelling and expansion of the hydrophilic push layer occurs, which puts additional force on the system to ensure drug delivery and release via the orifice. Additional variants of this “push-pull” system are also used in practice, such as the longitudinal compressed tablet (LCT) multilayer formulation, which can be adapted to contain multiple layers of drugs. The use of multiple drug layers provides the option

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Fig. 5 Illustratory diagrams of the drug release mechanism from (A) elementary osmotic pump and (B) osmotic push-pull system

of moving away from a zero-order release profile (which is achieved using the conventional push-pull system). For example, by formulating two drug layers with differing drug contents enables the modulation of the release rate profile. 2.1.4 Ion exchange-based formulations Ion exchange-controlled release systems comprise water-insoluble polymers functionalized with ionic groups that are capable of exchanging ions with those in a bulk solution with which it is in contact [4]. These systems involve drug molecules attaching onto the ionic moieties via electrostatic interactions due to opposite charges. Upon immersion into an aqueous environment, the drug molecules are replaced with ions with the same charge, in turn being released from the ion-exchange resin system (Fig. 6). In order to have a greater control (and achieve a more predictable drug release), the ion-exchange resins can be coated with water-insoluble polymers such as ethylcellulose, additionally providing a diffusion-controlled drug release. Overall, the rate of drug release depends on the area of diffusion (i.e., surface area of resin particles), cross-linked density, electrolyte concentrations in aqueous media (i.e., Na + or K+ for cationic drugs and Cl for anionic drugs), and coating thickness of the drug-resin complex.

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Fig. 6 Mechanism of drug release from ion-exchange-controlled drug release systems.

Pennkinetic systems enable an improved control over the drug-release rate (originally patented by Pennwalt Corp.) [5]. In these systems, the drug-resin complexes are pretreated with polyethylene glycol 400 (PEG 400), followed by subsequent coating with a water-insoluble polymer. PEG 400 aids in controlling the rate of swelling of the matrix in water, while the rate-controlling coating modifies the diffusion pattern of ions. Two over-the-counter (OTC) products, (dextromethorphan cough syrup and codeine and chlorpheniramine syrup) are examples of marketed formulations of Pennkinetic systems. Ion-exchange controlled release systems provide unique advantages for drug delivery. First, ion-exchange resins have a better capability to retain drug in the structure, thereby reducing the risk of dose dumping, which is a common challenge that extended release drug delivery systems face. Second, the resins can be formulated in a number of ways; as liquid suspensions, filled into capsules or compressed into tablets. Finally, due to the release of the drug in small quantities, there is a lower potential for local GI irritation, demonstrated in the case of diclofenac for the treatment of arthritic patients.

2.2 Delayed release formulations Delayed drug release is commonly achieved via enteric coating of dosage forms such as tablets, capsules and multiparticulates [6, 7]. Typically, enteric coatings are pH activated, wherein they are insoluble at low pH but dissolved readily at higher pH (e.g., pH 5–7) [8]. The main function of an enteric coating is to protect the underlying dosage form and drug substance, enabling it to remain intact the gastric environment and instead dissolve and undergo drug release in the small intestine [9, 10]. Such strategies are used to prevent gastric mucosa irritation caused by certain drugs (e.g., non-steroidal anti-inflammatory

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drugs; NSAIDs), or to avoid the degradation of acid-labile drugs in GI fluid, such as enzymes or peptides [11]. Protection can be easily and readily provided with the application of polymeric coatings that are inherently insoluble at acidic pH values. Enteric coatings can also be applied for the local treatment of intestinal diseases. For example, duodenal peptic ulcers are commonly treated locally against Helicobacter pylori with antibiotics like clarithromycin and amoxicillin in combination with acid blockers such as cimetidine or ranitidine. Medicines used to treat inflammatory bowel diseases (e.g., budesonide) also use delayed-release coatings using polymers to enable targeted in specific regions in the GI tract. However, several studies have shown that enteric-coated products designed to release in the proximal small intestine do not disintegrate rapidly after emptying from the stomach [12]. Indeed, it was shown in vivo that such products can take up to 2 h to disintegrate in the human small intestine [13–15]. Drug release will then occur in the distal small intestine and cause a delayed response to medication, and potentially reduce the bioavailability of those drugs with an absorption window in the upper small intestine. To overcome this challenge, and to ensure drug release at the upper small intestine, a novel double-layer technology marketed under the brand name DuoCoat was developed (Fig. 7). The outer layer of the DuoCoat technology is composed of a regular EUDRAGIT L 30 D-55 enteric coating that protects the dosage form during upper gastric transit and starts to dissolve at the pH of the upper small intestine (pH 5.5). The inner layer is a modified EUDRAGIT L 30 D-55 coating which has been neutralized by the addition of sodium hydroxide [16]. When the DuoCoat formulation enters the duodenum, the environmental pH value increases and at pH 5.5 the outer EUDRAGIT coating starts to swell and dissolve [17]. Intestinal fluid then penetrates into the system and reaches the neutralized inner coating layer, causing rapid dissolution of the coating and hence drug release [7]. Results from both in vitro biorelevant dissolution tests and in vivo human gamma scintigraphy studies showed that the the DuoCoat formulations released over twice as fast compared with a standard enteric polymer and demonstrated less variable results (Fig. 8) [7, 18].

Fig. 7 Diagram of the DuoCoat technology for targeted drug release in the upper small intestine.

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Fig. 8 In vivo gamma scintigraphy studies in eight human subjects of EUDRAGIT L30 D55 single coating vs. DuoCoat double coating. Reprinted with permission Liu F, Basit AW. A paradigm shift in enteric coating: achieving rapid release in the proximal small intestine of man. J Control Release 2010;147(2):242–245. https://doi.org/10.1016/j.jconrel.2010.07.105.

2.3 Targeted release formulations 2.3.1 Gastroretentive (GR) devices Provided that the drug is stable in an acidic environment, gastroretentive devices enable the delivery of drugs for an extended period of time in the stomach and upper GI regions compared with conventional dosage forms. In general, the retention of oral dosage forms within the stomach or upper GI tract regions is an attractive concept for those drugs with one or more of the following characteristics [19–21]: 1. Drugs that are locally active within the stomach (such as antacids and ulcer-treating drugs). 2. Drugs with a “window of absorption” within the stomach or upper GI regions (such as levodopa). 3. Drugs with high solubility at a low pH (such as chlordiazepoxide and diazepam). 4. Drugs with inherent instability within the lower GI tract regions and the colon (e.g., amoxicillin trihydrate). Thus far, several approaches have been used to attempt the long-term retention of dosage forms in the stomach, such as expanding devices; magnetic systems; floating systems; sinking systems and mucoadhesive systems (Table 1) [22–27]. It is worthy to note that despite the extensive research on GR devices, to date only a limited number of GR devices have reached the clinic. This is due to the systems having a number of disadvantages preventing their use clinically. For example, magnetic systems are highly dependent upon the position of the outer magnet for device success, and in most cases, it cannot be accurately located or affixed in place. Furthermore, this system

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Table 1 Examples of gastroretentive (GR) systems. GR system type

Mechanism

Expanding devices

An in vivo size change occurs, enabling the dosage form to exceed the diameter of the pyloric sphincter (either via swelling when in contact with GI fluid, or by unfolding in the stomach region), preventing dosage form emptying A magnet is embedded into the dosage form and another is placed on the outer abdominal region. The magnetic force created between the pieces holds the device in its place The dosage form has a lower density than the gastric fluid, enabling it to float on its surface. Initiation of floating is generated either via (1) generation of a gas due to the presence of an effervescent material or (2) use of a swellable polymer that traps air bubbles following its expansion Via the integration of heavy inert metals, the dosage form has a higher density than the gastric fluid, causing it to sediment in the stomach The system uses a bioadhesive polymer that adheres to the epithelial lining of the stomach to cause gastroretention

Magnetic systems

Floating systems

Sinking systems Mucoadhesive systems

is not well tolerated by patients as it causes discomfort. The floating and sinking devices are highly affected by the presence or absence of gastric contents, and exhibit vast interand intra-individual variability. The mucoadhesive systems can be accidentally removed due to the GI contractions, mucus turnover or due to the presence of food, leading to the premature ejection of the device. The vast majority of GR devices that have reached the clinic are reliant of the expanding device mechanism. An example of such is Depomed’s Acuform technology, composed of a combination of two polymers (polyethylene oxide; PEO and hypromellose methylcellulose; HMPC). When in contact with the gastric fluid, the dosage form undergoes swelling (PEO effect) and an extended drug release of up to 10 h (HPMC effect). To date, there are four marketed products based on this technology, including Janumet XR (sitagliptin and metformin), Glumetza (metformin), Gralise (gabapentin) and Nucynta ER (tapentadol). Another noteworthy technology is Intec’s Accordion Pill, which is reliant on an unfolding mechanism for gastroretention. The device is composed of a multi-layer film with varying release rates, which is folded in the shape of an accordion and is encapsulated for ease of oral administration. In the stomach, the capsule shell dissolves releasing the device, which then unfolds back to its original shape. Studies have shown that this system has a gastric residence time up to 12 h, providing a constant drug release during that period. This technology however, is still undergoing clinical trials and its action in vivo is yet to be proven. The most recently developed unfoldable GR device was developed by researchers at the Massachusetts Institute of Technology (MIT). The researchers developed a “starshaped” oral dosage form that can be retained in the stomach for seven days, allowing

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once-weekly dosing for antiretroviral drugs (dolutegravir, cabotegravir and rilpivirine) for HIV treatment [28]. From the outside, the drug product looks like a normal capsule. However, when it reaches the stomach, the capsule dissolves and a six-armed star-shaped structure that has been packed inside the capsule unfolds. Due to its geometry, prolonged gastric residence can be attained, as when expanded, the star-shaped structure is too large to move out of the pyloric sphincter of the stomach. APIs can be loaded into the arms of the star-shaped structure providing a prolonged release of the drugs over time, which will then degrade after drug release is completed and pass through the digestive tract [27]. 2.3.2 Colon-targeted drug delivery systems Colon-specific drug delivery has gained precedence in recent years not only for local drug delivery (i.e., in the treatment of Crohn’s disease, ulcerative colitis or colon cancer) but also for the oral delivery of drugs (such as proteins and peptides) that have been found to exhibit maximal systemic absorption from the colon. Critically, a successful colonspecific drug delivery system requires the provision of adequate protection for the drug en route to the colon i.e., by protecting it from release and degradation in the upper GI tract. Indeed, the system is required to release the drug at the colon that has close to neutral pH (7.5). The colon has been considered suitable for absorption of complex biomolecules due to the following advantages compared with other regions of the GI tract: 1. There is less prevalence and diversity of digestive enzymes, known to play a major role in the degradation of biomolecules in the stomach and upper intestinal regions. 2. Proteolytic activity of the colonic mucosa is significantly less compared with that of the upper GI tract mucosal regions. 3. There is a long gastric retention time for food and drug products in the colon (up to 5 days). To date, several approaches have been used in the development of colon-specific delivery systems, ranging from pH-controlled systems, time-controlled systems, microbiallycontrolled systems, and luminal pressure-controlled systems [29, 30]. The most traditionally used system are the pH-modulated systems, involving coating a dosage form (i.e., pellets or tablet core) with an enteric polymer, designed to enable drug release only in the higher pH of the colon [31–35]. However, there is growing evidence that these systems are not always consistent due to inter- and intra-individual variability in gastric pH and composition, gastric enzyme content, diet, varying transit time in the GI tract and disease state, amongst others [5, 36–40]. Indeed, several studies have highlighted how disintegration of the enteric polymer in the colon does not occur, thereby causing no drug release and hence therapeutic effect. To overcome this challenge, academic research has pursued the development of novel drug delivery systems that enable a better control of drug delivery and release at the colonic site. As an example, the world’s first dual-trigger colonic delivery system novel (Phloral) has been developed, comprising a pH-sensitive polymer and a natural polysaccharide that coats the dosage form that is digested by the colonic microbiota [41, 42].

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Fig. 9 Drug release of Phloral vs. conventional pH sensitive coatings. Circles represent the site of disintegration in individual subjects. Conventional dosage forms remained intact in subjects 4, 5 and 7 and passed out in the stools. Credit: IntractPharma. Phloral 2019 [23rd Oct 2019]. Available from: https://www.intractpharma.com.

The system exploits both the changes in gastrointestinal pH in combination with the enzymatic activity of the microbiota as independent but complementary release mechanisms to facilitate site-specific release. This technology provides a fail-safe mechanism, overcoming the limitations of conventional polymer coatings (Fig. 9). Following success in Phase III clinical trials, this innovative drug delivery technology has recently been launched as ASACOL 1600 mg (mesalazine) tablets across Europe, enabling delivery of the highest dose of any oral product in the world (1.6 g of drug) in a single dosage form, thereby reducing the frequency of administration to ensure patient compliance and therapeutic efficacy. It is likely that a similar concept could be applied for the delivery of peptides and large biomolecules (such as proteins and vaccine components) via the oral route. Due to the reduced presence of digestive enzymes and harsh pH conditions, the bioavailability of such biomolecules could be improved.

3. Novel technologies for modified drug release In recent years, novel modified release systems have been developed in order to orally deliver particularly challenging molecules (such as macromolecules, or for those most highly absorbing in the stomach region) to specific regions of the GI tract. Within the past decade, the use of novel technology and drug delivery techniques have been explored to achieve site-specific oral delivery systems, ranging from three-dimensional (3D) printing and even robotic microneedle pills for targeted GI delivery.

3.1 Microneedle pills for oral drug delivery In an attempt to overcome the harsh conditions of the GI tract, several research groups have explored the use of robotic pills. One such dosage form is the RaniPill (developed by Rani Therapeutics), which injects drugs directly into the small intestine (Fig. 10) [43].

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Fig. 10 Graphical illustration showing the drug release mechanism from the RaniPill. Credit: RaniTherapeutics. RaniPill: Our Technology 2019 [23rd Oct 2019]. Available from: https://www.ranitherapeutics.com/technology/.

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The robotic pill is a sophisticated device whereby the API is contained inside hollow and dissolvable microneedles, which is then assembled into a “balloon-like” structure and folded into the capsule body. The capsule has a pH-sensitive enteric coating, enabling it to pass through the stomach without being degraded. Upon reaching the upper small intestine (duodenum), the rise in pH causes the enteric coating to dissolve, initiating a chemical reaction to produce gas inside the balloon. Balloon inflation is triggered when citric acid reacts with sodium bicarbonate and this reaction creates the pressure needed to push the drug-loaded needles into the intestinal wall. Favorably, the intestinal wall is devoid of pain receptors, thereby enabling a pain-free alternative delivery of injectable drugs, whilst enabling an increased permeation of drugs across the GI mucosa. The technology could potentially be applied to solving oral delivery challenges for large proteins and macromolecules that typically do not permeate the GI mucosa. A similar concept has been developed by researchers at MIT in collaboration with Harvard Medical School, Brigham and Women’s Hospital and Novo Nordisk in an attempt to deliver insulin via the oral route. The team developed a device termed SOMA (Self-Orienting Millimeter-Scale Applicator) which was inspired by the by the selforienting leopard tortoise, a species that can flip itself over when on its back [44]. The SOMA’s shape and density distribution were optimized such that the microneedle lands in the stomach in the same orientation every time, enabling injection through the gastric epithelium.

3.2 Three-dimensional (3D) printing medicines To date, 3D printing has been used to create a range of complex formulations that would not easily be produced by conventional manufacturing technologies [45, 46]. This technology provides a high flexibility enabling the production of a multitude of drug products with tailored release profiles and designs, ranging from controlled-release formulations, fast-dissolving tablets and multi-drug combinations [46a,47–57]. To date, one 3D printed oral dosage form (Spritam) has been approved by the Food and Drug Administration (FDA) to treat epilepsy [58]. Furthermore, a world first clinical study has recently been conducted whereby a 3D printer was integrated into a hospital pharmacy to produce 3D printed medicines with a tailored dosage to treat pediatrics with a rare metabolic disease [59]. Drug release can be controlled by varying three main parameters in 3D printing; namely the printlet geometry, infill percentage and polymer inclusion. As an example, several studies have highlighted the ability for drug release to be tailored based on printlet design [60–63]. One study found that changing the surface area to volume ratios of different shaped printlets (cube, pyramid, cylinder, sphere and torus) enabled the modification of the drug release (Fig. 11) [64]. When tablets were formulated with a constant surface area/volume ratio, the release rates (fastest first) were in the order of sphere and cube>torus >cylinder >pyramid.

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r R h

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Fig. 11 Computer aided designs (top) and 3D printed tablets (bottom) of different geometries. Reprinted with permission from Goyanes A, Robles Martinez P, Buanz A, et al. Effect of geometry on drug release from 3D printed tablets. Int J Pharm. 2015;494(2):657–663. https://doi.org/10.1016/ j.ijpharm.2015.04.069.

Theophylline-loaded printlets with innovative “radiator-like designs” have also recently been developed [62]. Each dosage form had connected paralleled plates with inter-plate spacing of either 0.5, 1, 1.5 or 2 mm. The researchers found that the minimal spacing between parallel plates of the design should be 1 mm to enable an immediate drug release from the structures. Another study demonstrated the potential for printlets to be manufactured with cylindrical and gyroid lattice structures and demonstrated the ability to achieve customizable release characteristics based on the geometry selected, with lattice structures demonstrating faster drug release compared with the cylindrical tablet (Fig. 12) [65]. A recent study has demonstrated the potential for 3D printing pellets (miniprintlets) containing two APIs with differing release profiles [63]. Infill percentage (that is the degree to which the internal space will be filled from 0%, hollow, to 100%, solid), has also been found to be another determinant influencing the drug release [66]. Previous studies have shown that printlets with a lower infill percentage exhibit a faster drug release, whereas tablets with higher infill percentages showed extended release profiles (Fig. 13) [67]. On the contrary, in a study carried out by Chai et al., a change in infill percentage was exploited to create gastroretentive tablets [68]. This was mainly due to the difference in densities, wherein, tablets having 0–20% infill had a density that was lower than that of the fluid media, causing them to float. The buoyancy effect increased the residence time of the tablets in the gastric region, promoting drug absorption from the early part of the small intestine. However, such phenomenon is highly dependent upon a patient’s diet and thus, a high variability in performance is expected. By selecting the appropriate excipients and printing parameters, printlets (3D printed tablets) can be developed to have defined drug release profiles (e.g., by using different

PRINT ANALYSIS

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Fig. 12 The impact of 3D printed cylindrical and geometric lattice structures on drug release. Reprinted with permission from Fina F, Goyanes A, Madla CM, et al. 3D printing of drug-loaded gyroid lattices using selective laser sintering. Int J Pharm. 2018;547(1–2):44–52. https://doi.org/10.1016/j.ijpharm.2018.05.044.

Fig. 13 Images of 3D printed tablets with different infill percentages (top) and the impact of infill percentage on drug release (bottom). Reprinted with permission from Goyanes A, Buanz ABM, Basit AW, et al. Fused-filament 3D printing (3DP) for fabrication of tablets. Int J Pharm. 2014;476(1–2): 88–92. https://doi.org/10.1016/j.ijpharm.2014.09.044.

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rate-controlling polymers). Several studies have evaluated this concept for single drugs and polypills. For example, one study described the production of enteric-coated caplets by printing a methacrylic polymer-based shell surrounding a drug core. In 2017, Goyanes et al. showed that delayed release tablets could instead be produced by incorporating the drug within an enteric polymer, negating the need for an outer enteric shell [47]. Moreover, novel strategies such as four-dimensional (4D) printing could provide excellent promise within the pharmaceutical sector, especially for the advancement of modified drug delivery. 4D printing involves the fabrication of 3D objects that evolve to a pre-determined manner under the influence of external stimuli. As the goal of 4D printed objects is to alter their shape and size from external stimulation, the printed material (otherwise known as smart material) used is required to be responsive to stimuli varying from heat, pH, magnetic field or light.

4. Conclusions and future perspectives Modified release formulations have been widely adopted since the mid-1900s enabling a tighter spatial and temporal control over drug release compared with IR preparations, in turn prolonging and maintaining therapeutic effect to improve patient outcomes. Over the past few decades, significant technological advances have spurred the development of novel and targeted MR strategies, showing promise for the oral delivery of macromolecules, a concept that was previously considered impossible. Such innovations are forecast to revolutionize treatments for patients around the world by avoiding the need for highly invasive and costly injections, paving the way for the next generation of MR dosage forms.

References [1] Wang Z, Shmeis RA. Dissolution controlled drug delivery systems. In: Design of controlled release drug delivery systems. United States: McGraw-Hill; 2006. p. 139–72. [2] Reynolds TD, Mitchell SA, Balwinski KM. Investigation of the effect of tablet surface area/volume on drug release from hydroxypropylmethylcellulose controlled-release matrix tablets. Drug Dev Ind Pharm 2002;28(4):457–66. https://doi.org/10.1081/ddc-120003007. PubMed PMID: 12056539; eng. [3] Verma RK, Krishna DM, Garg S. Formulation aspects in the development of osmotically controlled oral drug delivery systems. J Control Release 2002;79(1-3):7–27. [4] Puttewar T, Kshirsagar M, Chandewar A, et al. Formulation and evaluation of orodispersible tablet of taste masked doxylamine succinate using ion exchange resin. J King Saud Univ Sci 2010;22(4):229–40. [5] Abuhelwa AY, Foster DJ, Upton RN. A quantitative review and meta-models of the variability and factors affecting oral drug absorption—part I: gastrointestinal pH. AAPS J 2016;18(5):1309–21. [6] Mehta LS, Gowda D, Gupta NV, et al. Formulation and development of lenalidomide loaded delayed release mini tablets in capsules. Int J App Pharm 2018;10(5):239–42. [7] Liu F, Basit AW. A paradigm shift in enteric coating: achieving rapid release in the proximal small intestine of man. J Control Release 2010;147(2):242–5. https://doi.org/10.1016/j.jconrel.2010.07.105. [8] Agyilirah GA, Banker GS. Polymers for enteric coating applications. In: Polymers for controlled drug delivery., vol. 3; 1991. p. 39–66.

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Delivery platforms for oral drug administration Katia P. Seremetaa,b, Alejandro Sosnikc a

Department of Basic and Applied Sciences, National University of the Chaco Austral, Pcia. Roque Sa´enz Pen˜a, Chaco, Argentina b National Scientific and Technical Research Council (CONICET), Pcia. Roque Sa´enz Pen˜a, Chaco, Argentina c Laboratory of Pharmaceutical Nanomaterials Science, Department of Materials Science and Engineering, Technion-Israel Institute of Technology, Haifa, Israel

1. Introduction Oral delivery remains the preferred route of drug administration due to its non-invasive nature and improved patient compliance. However, poor solubility, stability and bioavailability of many drugs make achieving therapeutic levels via the gastrointestinal (GI) tract challenging [1, 2]. The absorption of any chemical entity administered by the oral route is a complex process that is influenced by various factors and events related both to the organism and the drug. Factors related to the organism include physiological properties as the diameter, length and surface area of GI tract portion in which the absorption takes place, pH, degradation in the GI fluid, gastric emptying, intestinal transit time and metabolism in the gut wall or the liver; while factors related to the drug include physicochemical properties such as aqueous solubility and permeability across the intestinal epithelium that can limit the fraction of absorbed dose [3]. According to the Biopharmaceutics Classification System (BCS), active compounds for oral administration are classified into four classes based on high/low solubility and high/low permeability [4, 5]. Approximately 50–70% of drugs are poorly water-soluble and categorized as class II (low solubility-high permeability) or class IV (low solubilitylow permeability) [6–8]. This percentage increases for new chemical entities under different stages of development. In these cases, the aqueous solubility and membrane permeability are the limiting step and hinder the oral absorption and bioavailability of the drug [9]. Generally, a compound should first be dissolved in the GI fluid to reach to its target site. Therefore, several strategies are used in order to increase the solubility of poorly water-soluble drugs, and to enhance their absorption and bioavailability after oral administration [10], including, for example, the use of solid dispersions [11, 12], nanoparticles (NPs) [13, 14], nanocrystals [15, 16], microparticles (MPs) [17, 18], and inclusion complexes [19, 20].

Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00010-7

© 2020 Elsevier Inc. All rights reserved.

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In addition, many of the biological therapeutics such as proteins and peptides have poor oral bioavailability due to formidable barriers posed by the GI tract. Therefore, the oral administration of such compounds requires effective drug delivery systems to achieve improved bioavailability and successful use in humans. The design of these systems requires an understanding of intestinal physiology and the mucosal microenvironment [2]. Leal et al. [21] reviewed in detail the composition and physicochemical properties of the mucus, such as pore size, viscoelasticity, pH, and ionic strength of different organs, including those of the GI tract. The GI tract acts as a selective barrier for the transport of molecules, allowing the passage of nutrients and other molecules to the systemic circulation, and hindering the diffusion of pathogens, drugs, and macromolecules [21]. The mucus gel layer is a complex aqueous mixture of glycoproteins, lipids and salts that cover epithelial surfaces along the GI tract and acts as a barrier against xenobiotics and pathogens, but it also hinders the absorption of drugs and drug carrier systems such as MPs and NPs [22, 23]. To overcome this barrier, one of the factors that must be taken into account is the surface charge of the drug carriers. The mucus has anionic substructures that supply negative net charge. Therefore, uncharged and negatively-charged carriers can move across the mucus, but positively-charged ones usually cannot permeate through it, as they become immobilized by ionic interactions. However, when negatively-charged carriers reach the intestinal epithelium, cell uptake via endocytosis is less pronounced than for positively-charged counterparts. Thus, a strategy to overcome this obstacle could be the development of drug carriers that can change their zetapotential from negative to positive as they permeate mucus [23, 24]. Other strategies based on the breakdown of the mucus could be problematic, as this layer has a protective function. Therefore, other strategies have emerged, including the use of carrier systems at nanometric scale capable of permeating the mucus without destroying it or only to a very limited extent and transiently. Mucus permeating systems can be active and passive [25]. Passive systems keep interactions with the mucus at a minimum due to a slippery surface, as poly(ethylene glycol) (PEG) coated systems [26–28] and virus mimicking carriers [29]. Oppositely, active systems interact with the mucus by means of chemical reactions. These systems are based on carriers bearing free thiol groups and on proteolytic enzymes that are able to cleave disulfide bonds and peptide bonds of mucus glycoproteins, respectively [30]. In addition, active systems based on thiolated polymers (coined as Thiomers®) [31, 32] and zeta-potential change [33, 34] that avoid a back-diffusion of the particles out of the mucus layer are also a promising strategy [30]. Most licensed vaccines are injectable despite the majority of human pathogens infecting the body through mucosal sites. This is due to the fact that oral vaccines face many hurdles, such as the harsh gastric conditions. Therefore, the design and development of delivery systems that incorporate appropriate immune-stimulatory adjuvants and protect them from degradation in the stomach and intestine, increasing the uptake in the lymphoid tissue of the GI tract, are urgently needed [35–37]. Davitt and Lavelle [36] reviewed the opportunities, challenges and potential delivery solutions to facilitate the development of improved

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oral vaccines for infectious enteric diseases. Strategies as carrier systems including MPs and NPs, and lipid or enteric capsules were described [36]. Moreover, capsid viruses are a useful model for developing nanocarriers used as mucosal drug delivery systems, as they can diffuse through mucus as rapidly as through water, and penetrate across the epithelium, since they have a smaller size than the mucus mesh spacing and surfaces that do not stick to the mucus [38]. Several pH-sensitive nanocarriers have been designed to increase efficacy and reduce toxicity associated with off-target drug exposure [39]. For example, pH-dependent polymers such as methacrylic acid copolymers (e.g., Eudragit® L100) can be used for drug release in the intestine preventing chemical degradation in the stomach because these copolymers do not dissolve at gastric pH [40, 41]. One of the most promising methods for targeted delivery to cancer tissues of anticancer drugs or genes is based on receptor-mediated endocytosis of nanocarriers [42]. Cancer cells overexpress on their surfaces many tumor-specific receptor that exhibit high binding affinity towards specific ligands. For example, biotin (vitamin H or B7) is an essential vitamin for cell division and an important growth promoter for tumors. Biotin receptors are overexpressed in various tumor cell types more than other vitamin receptors. Therefore, biotin receptors can be used as biomarkers to design cytotoxic oral delivery systems modified on their surface with tumor-targeting ligands [43, 44]. Neonatal Fc receptor (FcRn) is related to the major histocompatibility complex Class I and interacts with antibodies of the immunoglobulin G (IgG) class. FcRn is expressed in several organs and tissues, in which it may have a role in IgG transport and protect IgG from degradation, extending its half-life in serum. Therefore, the conjugation of the IgG Fc portion to a nanocarrier can enhance the intestinal absorption and extend the half-life in the serum, improving the oral bioavailability and increasing tissue accumulation of the therapeutic antibody [45–47]. In this context, the administration frequency or dosage requirements could be reduced [48]. Sockolosky and Szoka [49] reviewed the therapeutic opportunities and challenges of targeting FcRn for drug delivery. Recently, FcRnmediated transcytosis has been proposed as a strategy to increase the transport of drugs across the intestinal epithelium [50].

2. Delivery platforms for oral administration of (bio)pharmaceuticals Different types of drug carriers have emerged as an effective strategy for mucosa delivery. One of the greatest challenges, however, is their ability to penetrate quickly through the mucus layer and this limits their success [51]. In this context, different modifications such as size and surface of the drug carriers can be introduced in order to achieve a greater uptake by the mucosa of the GI tract, and to enhance the bioavailability and therapeutic effect of orally administered drugs.

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2.1 Effect of the size and shape of the drug delivery system Controlling the size of the drug carriers is very important as it can determine changes in the pharmacokinetics and/or pharmacodynamics of the drugs they transport [52]. Moreover, drug loaded NPs, due to the reduction of the size and increase of the surface area, offer a versatile delivery platform to enhance the dissolution rate and bioavailability of poorly water-soluble drugs [53]. Ahmed et al. [54] prepared atorvastatin calcium (AC)-loaded NPs (AC-NPs) by the solvent-displacement method as an approach to improve the dissolution rate of AC. This is one drug among the statins that is widely used to inhibit cholesterol synthesis. However, the AC oral bioavailability is limited by its poor water solubility. For the preparation of the NPs, acetone containing poly-Ɛcaprolactone, the drug and Span® 80 were injected at controlled flow rate into an aqueous phase under constant sonication. Two formulations containing either Pluronic® 68 and Tween® 60 as stabilizers in the aqueous phase were obtained, and designated as FP-NPs and FT-NPs, respectively. After, acetone and water elimination by rotary evaporation, the recovered particles exhibited small mean particle size (200 nm) and span value (>1.2) exhibited lower blood bioavailability than Lipitor®, but higher efficacy to control hyperlipidemia with no adverse effects [55]. Thus, Cmax of these NPs was 98–135 ng/mL, while for Lipitor® was 285 ng/mL. In addition, AUC0–24 for the NPs and Lipitor® was 408–567 and 1214 ng h/mL, respectively. This indicates that these NPs retain the efficacy at lower doses, despite the lower relative bioavailability with respect to Lipitor®, due to higher drug accumulation in the liver, resulting in lower apparent bioavailability in blood, but better liver targeting. These results suggest that particle size plays an important role not only in blood bioavailability, but also in the pharmacological response [54, 55]. Small particles (about 50 nm) can easily go through epithelia by the paracellular pathway. Nevertheless, very small particles ( 200 nm > 500 nm > 1000 nm), and the uptake of all the particles was consistently higher in Caco-2/HT-29 cells compared to Caco-2 monolayers. In addition, unmodified spheres of sizes in the same range were incubated with a Caco-2 monolayer and co-cultures of Caco-2/Raji-B, Caco-2/HT-29 and Caco-2/HT-29/Raji-B cells, to evaluate the effect of the particle size on the transport across these models of the intestinal epithelium. Results showed that the transport across the Caco-2 monolayer was minimal and did not depend on the size. A maximum transport (0.3%) was observed for 50 nm sized particles after 5 h. The highest transport across Caco-2/Raji-B cells was observed for 50 nm sized particles (25%) and 200 nm (22%) after 5 h, and the transport across Caco-2/HT-29 cells showed similar profile to that of Caco-2 monolayers. In the triple co-culture, particles of 50 and 200 nm showed significantly higher transport (>20%) at 5 h compared to particles of 500 and 1000 nm (15% and 8%, respectively). These results indicate that Raji-B cells, which are derived from B lymphocytes, are essential for the transport of NPs across the semipermeable membranes. To evaluate the effect of the shape on the uptake and transport of NPs across intestinal cells, spheres, rods and discs were prepared. In Caco-2 cells, the uptake of rods and discs was significantly higher than that of spheres, but there was no difference in the uptake between rods and discs. However, in Caco-2/HT-29 cells, the uptake of rods was significantly higher than that of spheres and discs. Regarding the effect of particle shape on their transport across intestinal cells, no significant difference was evidenced on the transport of spheres, rods or discs in the Caco-2 monolayers. At the end of 5 h, a maximum transport of 0.5% of total particles was observed for all the different particles. In Caco-2/Raji-B co-culture, the transport of

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both rods and discs increased up to 20%, while that of spheres increased to a lesser extent (14%) after the same incubation time. In Caco-2/HT-29 cells, the particle transport did not show any difference among shapes. A maximum transport of about 1% was observed for spheres, but not significantly different from that of rods and discs after 5 h. The transport across the triple co-culture was similar to that of Caco-2/Raji-B. Discs showed the highest transport (18%) at the end of the assay, but not significantly different from that of rods (15%). On the other hand, spheres showed less transport capacity than that of rods and discs (11%). In summary, this study indicated that rods and discs are transported to a greater extent compared to spheres across intestinal cells. Confocal images of transwell membranes using the triple co-culture model at the end of the transport experiment were consistent with this observation. Finally, the effect of surface modification on the intestinal uptake and transport of NPs was evaluated using biotin-conjugated particles. Results showed that, regardless of the shape, conjugated particles are uptaken to a greater extent in both Caco-2 and Caco-2/HT-29 systems when compared to the unconjugated particles. Thus, the uptake of conjugated rods was 3- and 2.5-fold higher in both Caco-2 and Caco-2/HT-29 cells, respectively, compared to that of unconjugated rods. In addition, the uptake of biotin-conjugated rods in both Caco-2 and Caco-2/HT-29 cells was significantly higher than that of biotin-conjugated spheres and discs, while biotinconjugated discs showed significantly higher uptake than that of biotin-conjugated spheres. Therefore, biotin receptor-mediated targeting considerably improves cellular uptake in both Caco-2 and Caco-2/HT-29 cells of rods compared to their non-targeted counterpart, and conjugated/unconjugated spheres and discs. Confocal microscopy images further support these results (Fig. 2). On the other hand, results of the transport across monolayers showed that the transport of conjugated particles was higher across Caco-2/Raji-B cells. Rods were transported to the highest extent (27%), followed by discs (23%) and spheres (13%) after 5 h. In the triple co-culture model, by the end of the study, conjugated discs were transported to the highest extent (21%), followed by rods (18%) and spheres (12%). The percentage of transport of biotin-conjugated particles was similar to that of unconjugated particles, with the exception of conjugated rods that showed significantly higher transport across the triple co-culture compared to unconjugated rods. This may be due to the fact that Caco-2 cells, which display biotin receptors, are not primarily responsible for transcytosis of NPs across the cells and, therefore, surface modification considerably improved their uptake, but did not increase their transport [59]. The nanocarrier shape has emerged as an additional feature that can have a significant effect on drug dissolution, absorption and oral bioavailability and efficacy [61, 62]. Zhang et al. [62] evaluated the effect of the shape of mesoporous silica nanoparticle (MSNs) on oral delivery of indomethacin (IMC) used as a poorly water-soluble model drug. They synthesized mesoporous silica nanorods (MSNRs) and mesoporous silica nanospheres (MSNSs) loaded with IMC. In vitro release assays showed that IMC was released faster

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Fig. 2 Confocal microscopy micrographs of the uptake of nanoparticles of different shapes and surface chemistries by Caco-2/HT-29 cells. Representative confocal microscopy images showing uptake of (A) spheres; (B) rods; (C) discs; (D) biotin conjugated spheres; (E) biotin-conjugated rods and (F) biotin-conjugated discs by Caco-2/HT-29 cells at the end of 5 h of study. Green fluorescence represents particles, while blue color represents cell nuclei (DAPI staining). Scale bar ¼ 10 μm. Adapted from Banerjee A, Qi J, Gogoi R, Wong J, Mitragotri S. Role of nanoparticle size, shape and surface chemistry in oral drug delivery. J Control Release 2016;238:176–85 with permission of Elsevier.

from MSNRs (100% within 1 h) than from MSNSs (about 80% within 1 h), which may be attributed to the different pore architectures of MSNs. Both samples increased the dissolution of IMC with respect to bulk IMC (27% dissolved after 1.5 h). The oral bioavailability study in male Sprague-Dawley rats showed that the AUC for IMC-MSNRs was 1.3-fold and 2.2-fold higher than that for IMC-MSNSs and IMC solution, respectively. Therefore, the bioavailability of IMC-MSNRs was 2.2-fold and 4.0-fold higher than that of IMC-MSNSs and IMC solution, respectively, which achieved the highest Cmax (928 mg/mL); while Cmax values for MSNSs and IMC solution were 417 and 234 mg/mL, respectively. These results indicated that MSNRs may be more effective than MSNSs for oral drug delivery [62] and are in agreement with those reported by Banerjee et al. [59]. In addition, Yu et al. [63] reported that MSNRs and calcium phosphate nanorods have superior transport and trafficking capability in mucus compared

Delivery platforms for oral drug administration

with spheres. This could be attributed to the rotational dynamics of the MSNRs facilitated by the mucin fibers and the shear flow. Most of the bacteria as Helicobacter pylori and Vibrio cholera have rodlike shape, which confers high mobility in mucus [63].

2.2 Effect of the surface of the drug delivery system The surface of the particles is also a parameter to take into account in oral administration due to its possible interactions with the mucus. Generally, carriers with negative charge are more slippery because they avoid ionic interactions with the mucus layer, which has a net negative charge and therefore, permeate through this layer more easily. However, when reaching the epithelium, negatively-charged carriers have a less pronounced cellular uptake via endocytosis than the positively-charged counterparts. Oppositely, carriers with positive surface charge or a less negative charge than the mucus tend to adhere to the cell membrane. In this sense, efforts have been devoted to the development of nanocarrier systems that change their zeta-potential in order to overcome both the mucus layer and the intestinal epithelium [64]. In order to achieve both strong mucus permeation and efficient epithelial absorption, Wang et al. [65] modified the surface of silica NPs with a cell penetrating peptide (CPP) layer (a polycationic peptide that interacts strongly with negatively charged proteins) and a succinylated casein (SCN) layer. The final system aimed to minimize the release of the loaded drug in the stomach and to improve the intestinal uptake of the nanocarriers after degradation of the SCN layer. Both paclitaxel (PTX) and quercetin (Qc), two extremely poorly water-soluble molecules, were used as model drugs. PTX has low solubility and permeability, which causes a very low oral bioavailability. In addition, this drug is extremely sensitive to the P glycoprotein (P-gp) efflux pump; therefore, Qc, a P-gp inhibitor, was co-loaded to improve the cellular uptake of PTX. The nanocarriers were prepared through a three-step approach. Quantum dots were encapsulated in hollow silica NPs (HSQN) to monitor the nanocarrier behavior and cellular trafficking in real time. First, CPP was conjugated to the surface of HSQN to prepare HSQN@CPP. Then, the positively charged HSQN@CPP were coated with the negatively-charged SCN to form HSQN@CPP@SCN nanocarriers. The final carrier was collected by centrifugation. Both PTX and Qc were loaded into HSQN@CPP@SCN. The average diameter of the carriers was approximately 180 nm. PTX was loaded most efficiently into the core via incubation (loading efficacy as high as 50%), while Qc, loaded after PTX, was retained in the outer wall (loading efficacy, 8–10%). The stability study in gastric and intestinal media showed that HSQN@CPP@SCN were stable at low pH, while the SCN layer gradually degraded under neutral conditions. The zeta-potential of HSQN@CPP@SCN changed from negative, when in simulated gastric juice, to positive in the intestinal one with trypsin, indicating that the SCN was degraded from the surface of the nanocarriers. In gastric juice, less than 10% of the drug was released during

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the first 2 h and, in the intestinal one, more than 50% of the drug was released over 24 h under sink conditions. This demonstrated that nanocarriers with a protective SCN layer are resistant to the low pH and pepsin, and that they can prevent the loss of drug in the stomach. Thus, the addition of the SCN coating localizes the release in the small intestine and minimizes the side-effects. Finally, this SCN layer can be disassembled by the intestine trypsin, exposing the CPP layer for transepithelium transport. The mucus permeation study was carried by incubating particles in mucin. The motion of the carriers in mucin solutions was observed by means of realtime multiple-particle tracking analysis. HSQN@CPP@SCN exhibited improved motility compared with HSQN@CPP due to the SCN layer, allowing for a better permeation of the HSQN@CPP@SCN. In addition, the fluorescence of the supernatants was measured after the nanocarriers were mixed with mucin solution and the suspensions were centrifuged after different incubation time. The fluorescence intensity decreased dramatically when CPP was conjugated to HSQN, which indicated strong binding affinity with the mucin. When the layer of SCN was added, most formulations remained unbound, and the fluorescence intensity of the supernatants was strong. A study of the interaction between HSQN@CPP@SCN treated with/without trypsin and cellular membranes was carried using the membrane probe (Nile red) method and quartz crystal microbalance. Thus, HSQN@CPP@SCN were incubated with Nile Red liposome (30 min) and the fluorescence spectrum of this membrane probe was scanned at λ ¼ 542 nm. Results showed that HSQN@CPP@SCN treated with trypsin interacted strongly with epithelial cell membranes, and that the disintegration of the SCN layer exposes the CPP layer to the intestinal epithelial cells, increasing the cell uptake and transepithelial transport of the nanocarriers. After incubation of trypsin-treated HSQN@CPP@SCN with liposomes, the fluorescence intensity decreased with respect to untreated NPs, which suggests that the degradation of SCN with trypsin induced a stronger interaction of the nanocarrier with liposomes. Results of quartz crystal microbalance analysis showed that very little binding occurred when liposomes or Caco-2 cells immobilized on the quartz crystal microbalance sensors were exposed to HSQN@CPP@SCN. However, the binding increased sharply after the nanocarrier was treated with trypsin, indicating that strong interactions were observed after the SCN layer was lost. Transport across Caco-2 cells showed that the uptake of nanocarriers was slightly improved after modification with CPP and SCN layers. When HSQN@CPP@SCN was treated with trypsin for 1 h, a larger fraction of nanocarriers was taken up by Caco-2 cells. The transport of carriers across the intestinal epithelium was studied in sections of duodenum, jejunum and ileum, 2 h after the oral administration of nanocarriers and animal disection. Image of confocal laser scanning microscopy showed that HSQN@CPP@SCN exhibited significantly more transepithelial transport than HSQN@CPP in all three intestinal segments, and strong fluorescence signal was detected for HSQN@CPP@SCN in the villi and epithelial cells. The oral bioavailability assay in rats showed that PTX and Qc co-loaded in HSQN@CPP@SCN led to higher

Delivery platforms for oral drug administration

plasma PTX concentration, with the maximum being reached at 3 h post-administration. Conversely, HSQN@CPP@SCN without Qc exhibited a relatively low PTX plasma concentration. In addition, HSQN@CPP@SCN loaded with PTX and Qc displayed an absolute bioavailability that was much higher (39.0%) than that of HSQN@CPP (11.6%) and the AUC0–1 of PTX after administration of HSQN@CPP@SCN was enhanced by 7.8-fold over oral Taxol (a commercial PTX medication administered through IV injection in the clinic, considered to have 100% bioavailability). An efficacy study carried out in mouse indicated that HSQN@CPP@SCN were more efficient in inhibiting tumor growth compared to HSQN@CPP, and that HSQN@CPP@SCN can induce greater tumor necrosis than the other groups. Results showed that the double modified HSQNs have a better anti-tumor effect; HSQN@CPP@SCN delivered twice as much PTX to the tumor than HSQN@CPP. In addition, the liver distribution of PTX in HSQN@CPP@SCN was higher than with HSQN@CPP. Mice treated with HSQN@CPP@SCN remained healthy with stable body weight throughout the experiment, but mice treated with HSQN@CPP showed signs of toxicity and inflammation in the different intestinal sections, likely caused by the premature leakage of cytotoxic drug. In summary, this is a novel oral drug nanocarrier to overcome the challenges of balancing mucus layer penetration and epithelial uptake by exploring surface modification strategies [65]. In another work, Inchaurraga et al. [66] prepared NPs from a copolymer of methyl vinyl ether and maleic anhydride “decorated” with PEG of different molecular weights (MW; 2000, 6000 and 10,000 g/mol), identified as NP2, NP6 and NP10, respectively, with the aim to evaluate their mucus-permeating properties following oral administration. The PEG-to-polymer ratio selected was of 0.125. Thus, NPs were obtained by a simple desolvation procedure after the incubation between the PEG and the polymer in acetone. All the NPs were spherical with mean size of approximately 150–160 nm and PDI < 0.2. Average sizes slightly increased by increasing the MW of PEG. The zetapotential of the NPs modified with PEG was slightly negative (44 mV) and less negative than that of the “naked” NPs (52 mV). The yield of the process was about 90%. Biodistribution studies were conducted with 99mTc-labeled NPs orally administered in male Wistar rats. Results showed that an important fraction of the “naked” NPs was retained in the stomach, whereas PEG-coated NPs were found mainly in the small intestine, at 3 h post-administration (Fig. 3). Studies of mucus-permeating properties of NPs were carried out in male Wistar rats using rhodamine B isothiocyanate-loaded NPs. Results showed that unmodified NPs interacted fast and to a higher extent with the gastric mucosa than PEG-coated counterparts. Regarding the PEG-coated NPs, the lower the MW of the PEG, the greater the adhesion. At 1 h post-administration, the amount of NP2 adhered to the stomach was at least 2-times greater than NP6 and NP10. For NP6 or NP10, the adhered fraction was always below 7% of the given dose. The profile of the curve for “naked” NPs showed that

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Fig. 3 Volume rendered fused SPECT-CT images from representative animals 3 h after administration of 99mTc-labeled NP by oral gavage. NP: “naked” nanoparticles; NP2: poly(anhydride) nanoparticles coated with PEG2000; NP6: poly(anhydride) nanoparticles coated with PEG6000; NP10: poly(anhydride) nanoparticles coated with PEG10000. Adapted from Inchaurraga L, Martín-Arbella N, Zabaleta V, Quincoces G, Peñuelas I, Irache JM. In vivo study of the mucus-permeating properties of PEG-coated nanoparticles following oral administration. Eur J Pharm Biopharm. 2015;97(Pt A):280–89 with permission of Elsevier.

the maximum of the curve (15%) was produced at 30 min post-administration and that the adhered fraction decreased rapidly. On the contrary, NP2 showed a profile characterized by a first step, in which the adhered fraction increased up to 22% 1 h postadministration, followed by a slow decline. NP6 and NP10 showed similar profiles but the adhered fractions of NPs increased slowly between 1 and 3 h. Then, the adhered fractions decreased rapidly. NP10 adhered significantly less than NP6 or NP2. In the small intestine, the AUCadh of NP6 was higher or equal to NP2, and the AUCadh of NP2 was higher than NP10 and “naked” NPs. Thus, the intensity of the adhesive interactions for NP2 and NP6 was 1.5- and 3-times higher than for NP10 and “naked” NPs, respectively. Moreover, the mean residence time of the adhered fraction for NP2 and NP6 was about 35 and 45 min longer than for NP10 and “naked” NPs, respectively. PEG-coated NPs were capable of reaching the epithelium and interact with the intestinal cells, being these interactions stronger for NP2 and NP6 than for NP10. Another aspect of consideration is that excessive PEG surface density tends to diminish the mucuspermeating properties and thus, both length and density need to be optimized [66].

Delivery platforms for oral drug administration

Corpas-Lo´pez et al. [67] obtained gold NPs by the sodium citrate reduction method as a promising oral nanocarrier for the treatment of visceral leishmaniasis, the most severe form of the disease caused by L. donovani and L. infantum (syn. L. chagasi), which proliferate as intracellular amastigotes in the bone marrow, the liver and the spleen. Gold NPs were coated with MDG (SAHA-OBn), a derivative of vorinostat or suberoylanilide hydroxamic acid (SAHA) belonging to the O-alkyl hydroxamate family without inhibitory activity on human histone deacetylase enzymes. SAHA is a hydroxamate with histone deacetylases (HDAC) pan-HDAC inhibitory activity that have moderate active against L. donovani. HDACs are involved in silencing critical regulatory pathways, including pro-apoptotic programs. Additionally, bovine serum albumin (BSA) was used as a protective polyelectrolyte/polymer to avoid NP aggregation. The evaluation of vorinostat, tubacin and valproic acid (inhibitors of mammalian HDAC) was also carried out against the parasite. However, none of these compounds was effective against L. infantum promastigotes at up to 100 μM and the viability remained 100%. They were not effective against intracellular amastigotes either. However, MDG displayed a high activity against L. infantum and L. donovani amastigotes without causing cytotoxicity on mammal cells at 100 μM, the highest concentration tested. Selectivity indices were higher than those of the reference compounds, suggesting a wide treatment window without inhibitory activity over mammal histone deacetylases. All evaluated concentrations of MDG (75–100 μM) were found to be not mutagenic, and this compound does not contain base-pair or frame-shift mutagens under the experimental conditions used on S. typhimurium TA98 and S. typhimurium TA100. Then, the vehiculization of MDG on gold NPs was carried out to enhance the pharmacokinetics and the efficacy and to reduce the toxicity. The average hydrodynamic diameter, determined by dynamic light scattering (DLS), for bare gold NPs was 21.5  2.5 nm and for MDG-coated gold was 58.6  6.5 nm, at pH 7. In turn, the diameter of the BSA-coated MDG-gold NPs was, for the different concentrations of albumin MDG-gold coated NPs, 142.7 and 134.3 nm, respectively. The surface charge of colloidal particles plays an important role in blood half-lives of drug delivery systems and in aqueous colloidal stability. In order to optimize the final surface charge best suited for drug delivery, bare gold NPs and MDG-coated ones were investigated by electrophoretic mobility. Gold NPs presented negative electrophoretic mobility over the whole range of pH. This negative charge is associated with the citrate shell. Higher concentrations of albumin produced a larger decrease of electrophoretic mobility. The color of the suspension remained reddish-pink after the blocking process due to the coating with MDG. NPs could easily be resuspended with a gentle shaking and, after several months, no color change was observed. A biphasic release profile was observed for the MDG adsorbed onto the surface over 3 days. In addition, MDG was stable at pH values similar to the stomach, which is relevant for oral administration. Anti-leishmanial activity and safety assays in the murine model of visceral leishmaniasis showed that MDG was effective at two dosage regimens (10 and 25 mg/kg) and through intraperitoneal and oral administration. The higher monotherapy regimen was more effective via intraperitoneal route in the spleen

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and liver. In the spleen, 25 mg/kg was more effective than the lower dose. On the other hand, the oral regimens were more effective at the lower dose in the spleen, and there was no statistically significant difference between the two MDG monotherapy doses employed in the spleen and liver. Meglumine antimoniate was efficacious only in the liver. The parasite burden reduction driven in the spleen or in the bone marrow was not statistically significant with this compound. However, when MDG (10 mg/kg) and MA were combined, MDG increased the efficacy of MA and the parasite load reduction was very high, particularly through the oral route. Thus, a parasite burden reduction of 95%, 95% and 99% in spleen, liver and bone marrow, respectively, was reached and parasite clearance was achieved in more than half of mice treated. The bone marrow is a relevant target tissue in the progression of visceral leishmaniasis and the parasite elimination is crucial for treatment success. The toxicity study on infected and healthy mice showed that the evolution of weight was similar in infected and untreated animals, without weight loss throughout the experiment. Both liver and spleen weights were similar among these groups. In addition, no signs of pain, stress, diarrhea and death was registered in treated groups, and there were no statistically significant differences in the transaminases, alkaline phosphatase, urea or creatinine values. The subacute toxicity study yielded the same results. In conclusion, the authors have shown that MDG can have a potent anti-leishmanial activity. The combination of MDG (10 mg/kg, oral route) and meglumine antimoniate resulted be the best treatment regime which decreased parasite load in the major target organs in more than 95%, achieving parasite clearance in more than half of mice [67]. These results are promising considering that, currently, there are no vaccines that protect against leishmaniasis [68]. Therefore, a potential antileishmanial drug should be effective, safe and affordable. In addition, it has to display high selectivity and potency against the parasite and to be able to be administered by the oral route with a high safety ratio in humans. However, current drugs present high toxicity, adverse effects, resistance, high cost, and geographical variability. In addition, they are expensive and have to be applied over a long period of time, mainly by injection [67, 68]. The transmembrane P-glycoprotein (P-gp) is found in the small intestine, the bloodbrain barrier and the surface of tumor cells. In the intestine, the P-gp can efflux the drug from the basolateral to the apical side of the epithelium. Therefore, the inhibition of P-gp at the intestinal level through carriers with modified surfaces may improve the oral bioavailability of drugs due to enhanced absorption [69]. When PEG is conjugated to the surface of a drug nanocarrier, it can improve the physical stability in the biological fluids and thus, to enhance their transport across mucosal surfaces [70]. Chitosan also can be used to alter the surface of drug nanocarriers and allow extended release, while conferring a mucoadhesive nature that enhances drug retention and absorption [71]. In summary, there are several strategies based on the modification of the carrier surface to increase permeability across the intestinal mucosa and therefore improve the absorption and bioavailability [72].

Delivery platforms for oral drug administration

2.3 Effect of the composition and the preparation method of the drug delivery system The adjustment of the process parameters (e.g., equipment configuration, flow rate, presion) and properties of the sample (e.g., type of solvent and polymer, viscosity, surface tension, addition of surfactant) allows the design of drug-loaded carriers with specific and controllable morphology for oral administration [73, 74]. The sample composition and preparation methods of these carriers can influence their behavior in vitro and in vivo. For example, Cirri et al. [75] prepared nanostructured lipid carriers (NLC) loaded with hydrochlorothiazide (HCT) to improve its therapeutic efficacy and obtain an oral pediatric formulation. HCT presents low solubility and permeability, with variable and limited bioavailability and there are not liquid formulations of this drug on the market. HCT-NLC were prepared by two different methods: homogenization-ultrasonication (HU) and microemulsion (ME). Transcutol® HP and Precirol® ATO5 (90:10 w/w ratio) were selected as liquid and solid lipid, respectively, and Tween® 80 or Pluronic® F68 were used as surfactants. The mean particle size of the NLC with Pluronic® F68 obtained by HU was smaller than those containing Tween® 80 and obtained by same method. An analogous result was observed for the entrapment efficiency and loading capacity. In addition, the type of surfactant influenced the drug release profile. Thus, NLC containing Pluronic® F68 showed a slowing down of the drug release rate, with approximately 20% drug released, while those with Tween® 80 exhibited a better drug release profile compared to the simple drug suspension (40%), with nearly 60% drug released at 120 min. However, due to that, a complete drug release was not achieved. Thus, new NLC were prepared by using mixtures of Tween® 80:Tween® 20 at 1:4 w/w ratio, Tween® 80: Solutol® HS at 1:1 w/w ratio, and Pluronic® F68:Tween® 20 at 1:2 and 1:4 w/w ratios by ME. Pluronic® F68 did not give rise to stable NLC. On the contrary, when Tween® 80 or Tween® 20 and Solutol® HS were used as co-surfactants, stable NLC were obtained, with particle sizes of approximately 330 and 430 nm, respectively, i.e., smaller than those of NLC obtained with the HU method. Furthermore, these NLC had higher encapsulation efficiency than the counterpart produced by HU method, with values of encapsulation efficiency ranging from 86% with Solutol® HS to 93% with Tween® 20, while only 80% were achieved by HU. Therefore, NLC obtained by ME were stable in buffer of pH 4.5 without any significant change in the mean size and PDI. In vitro release studies showed prolonged release, which lasted for 6 h, with both formulations achieving 80–100%. These results confirmed that the ME technique is the most effective method to produce HCT-loaded NLC. In vivo studies in normal adult male rats showed that the HCT-NLC formulation obtained by ME using Tween® 80 with Tween® 20 displayed the best diuretic profile, superior (P < 0.05) to that of HCT suspension at 4–6 h posttreatment. In addition, this formulation exhibited the highest diuretic index of all the groups treated with other formulations obtained by ME. The AUC value of NLC with Tween® 80/Tween® 20 was approximately 20% greater than that of the drug suspension.

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Stability studies of stored NLC showed no substantial changes of PDI values for all NLC series obtained by ME during the considered storage period of 3 months. The particle size of all formulations showed moderate reduction. Overall, these findings stress the key role of the components and the preparation method in the performance of the NLC [75]. Design of experiments (DoE) is a useful tool for formulation optimization. For example, Singh and Pai [76] obtained an optimized formulation of Eudragit RL100 NPs loaded with the HIV protease inhibitor atazanavir (ATV), after applying 32 central composite design. ATV displays low oral bioavailability due to poor water solubility, pH-dependent dissolution, effect of rapid first-pass metabolism and P-gp efflux. NPs were obtained by the nanoprecipitation method using acetone as organic solvent to dissolve the polymer and drug. This organic phase was dropped into aqueous phase with surfactant tetradecyl trimethyl ammonium bromide while homogenizing (12,000 g) during 30 min. Then, this emulsion was magnetically stirred (5 h) at room temperature to evaporate the organic solvent. Finally, NPs were centrifuged, washed and lyophilized. The yield obtained was 97.3%. The critical influential factors selected were the polymer and the surfactant. These factors were studied at three levels each by using 32 CCD. Particle size, drug encapsulation efficiency, and mean percentage drug release in 24 h were encompassed in the DoE. The optimized NPs were selected by overlay plot. The criteria of selection was: size 56:9% and mean percentage drug release in 24 h >85%. Results showed that the NP size was between 390 and 487 nm. Size was smaller at higher amount of surfactant and intermediate amount of polymer. Drug encapsulation efficiency values were between 41.3% and 56.9%. This indicates that the polymer used (Eudragit RL 100) and the nanoprecipitation technique can give high encapsulation efficiency for poorly water soluble drugs such as ATV. Regarding the mean percentage drug release in 24 h, in pH 7.4 PBS at 37 °C, all the NPs released a high percentage of the drug (85%) within 24 h and in a sustained manner, due to the used polymer. The optimized formulation contained 278 mg of polymer and 249 mg of emulsifier, and exhibited particle size, drug encapsulation efficiency, and mean percentage drug release in 24 h of 465.59 nm, 59.28% and 89.33%, respectively. In vivo pharmacokinetic assays in Wistar rats showed enhanced bioavailability of the optimized formulation with respect to the pure drug, with an increase of 1.10- and 2.91-fold in Cmax and AUC0–24, respectively. In situ single pass perfusion studies in Wistar rats showed that effective permeability of drug was 2.11-fold higher in optimized NPs with respect to ATV. In summary, optimization of ATV-loaded NPs by DoE is an efficient strategy to design and obtain a formulation with desired size, encapsulation efficiency and release, and then conduct the in vivo studies with promising results [76].

2.4 Targeted drug delivery systems An effective strategy for targeting the delivery of therapeutics agents by using nanocarriers, for example, could enhance the transport of active substances to specific target

Delivery platforms for oral drug administration

tissues and avoid or reduce the undesired side-effects caused by these agents in other organs or tissues [52]. Drug nanocarriers can be designed for targeting drugs to different portions of the GI tract [77]. In this sense, Hyun et al. [78] developed a nanocomplex with a hyaluronic acid-taurocholic acid (HA-TCA) conjugate vehicle for oral siRNA delivery with therapeutic potential in colorectal liver metastasis (Fig. 4). Protamine sulfate was used as vehicle for siRNA due to its cationic property. HA-TCA siRNA/protamine (sRP) nano-complexes (HTsRP-NCs) were obtained by a charge to charge complexation between HA-TCA and sRP. The increased molar ratio of protamine to siRNA led to either a decrease in the size or a change of the surface charge from negative to positive of the sRPs. An sRP complex with a molar ratio of 1:20 (sRP20) was selected and complexed with negatively charged HA-TCA to produce HTsRP-NCs. Various molar ratios of protamine to HA-TCA (1:2, 1:10, 1:20, 1:40)

(A) siRNA

Protamine

(D) Enterohepaic Circulation

(B) sRP-NCs

HA-TCA

CLM

(E)

(C) HTsRP-NCs

Fig. 4 Scheme of HTsRP-NC system for oral delivery of therapeutic siRNA. (A) Formation of sRP-NCs by a electrostatic complexation of siRNA and protamine, (B) formation of HTsRP-NCs by layering of the sRP-NCs complex with HA-TCA conjugates, (C) oral administration of the HTsRP-NCs, (D) TCAmediated Ileac absorption of the HTsRP-NCs from the intestinal apical region to the enterohepatic circulation system, (E) hyaluronic acid (HA) receptor-mediated selective uptake of HTsRPNCs to the colorectal liver metastasis (CLM). Adapted from Hyun EJ, Hasan MN, Kang SH, Cho S, Lee YK. Oral siRNA delivery using dual transporting systems to efficiently treat colorectal liver metastasis. Int J Pharm 2019;555:250–258 with permission of Elsevier.

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were used to prepare the complexes with 0.1 nmol of SiRNA. The higher the HA-TCA ratio, the greater the particle size and the less positive the surface charge. A molar ratio of 1:40 (HTsRP40-NCs) resulted in particle diameter of 246  6 nm and zeta-potential of 19.5  0.1 mV, and the siRNA encapsulation efficiency was 96.7%. siRNA oral delivery remains a challenge due to the harsh acidic pH of the stomach and the enzymatic degradation in the GI tract. Thus, the authors analyzed the stability of siRNA loaded in HTsRP-NCs in the simulated pH conditions of the GI tract (pH of 7.4, 5.6 and 1.2). Results of DLS analysis indicated no significant size changes for up to 12 h. Regarding enzymatic degradation, incubation of HTsRP40-NCs with either RNase or serum showed that siRNA was effectively protected in HTsRP40-NCs from RNase during 1 h of incubation. In turn, naked siRNAs were degraded within 30 min. In addition, siRNA in HTsRP40-NCs was retained for 12 h under 50% serum, while naked siRNA was rapidly degraded under the same condition. The release of siRNA was sustained, regardless of the pH values. However, the release rate at pH 5.6 and 7.4 was significantly increased with respect to pH 1.2, indicating the protection against the harsh acidic conditions. Additionally, in presence of hyaluronidase, the release of siRNA showed a dramatic increase (90%) during 24 h of incubation. This result suggests a potential use of HTsRPNCs to deliver siRNA in the targeting of tumors, since the cytosol of tumor cells has abundant hyaluronidase. The cellular cytotoxicity assay of the HA-TCA/protamine complex demonstrated very low cellular toxicity in MDCK cell line, with over 80% cell viability at a high concentration (12 μg/mL), suggesting its feasibility as a siRNA carrier. Furthermore, rhodamine B-labeled protamine/HA-TCA complex (HTrBP-NC) was assayed in HepG2 (liver cancer cells) and HCT-116 (colorectal cancer), two different CD44 positive cancer cell lines, with the aim of studying the potential delivery to the tumor site. A therapeutic target of siRNA is the HA receptor (CD44), overexpressed on the tumor cell surface. Results showed the presence of rhodamine B in both CD44positive cancer cells due to the interaction between CD44 and HA of the carrier. In addition, HTRP-NC was tested in MDCK-ASBT cells overexpressing ASBT receptors to investigate whether these carriers can use bile acid-mediated intestinal uptake. MDCK-ASBT cells exhibited about 3 times more cellular uptake when compared to MDCK cells (negative control) by measuring fluorescence intensities. Therefore, results suggest ASBT mediated HTsRP-NCs absorption in the small intestine, and the possible delivery to the liver through the enterohepatic circulation. The authors demonstrated that the protamine component complexed with siRNA in HTsRP-NCs is the cationic material that induces a “proton sponge effect” for successful escaping from the endosomal degradation of siRNA by the increase of the osmotic pressure in the endosome. This phenomenon is known as “endosomal escape”. A biodistribution assay in BalB C mice after oral administration of HTsRP-NC containing the near infrared dye NIR-670 instead of siRNA indicated a significant accumulation into liver and small intestine, suggesting that the carriers can be delivered from the small intestine to the

Delivery platforms for oral drug administration

liver through the enterohepatic portal system. Subsequently, HTsRP-NCs containing AKT (protein kinase B) siRNA were orally administered to a colorectal liver metastasis murine model. An almost non-detectable tumor signal in the mouse treated with HTsRP-NCs at day 16 was observed by optical imaging when compared to the control tumor murine model. In addition, a reduction of the total number of liver tumor nodules larger than 1 mm in size was found in mice treated with HTsRP-NCs. Finally, the reduction of pAKT expression and the activation of Bax, an apoptotic protein, after immunoblot analysis of liver tissue obtained from the mice treated with HTsRPNC was evident. This points out the potential of siAKT-RNA HTsRP-NC in the treatment of the colorectal liver metastasis in this murine model by the oral route [78]. Rectal route has drawbacks such as limited treatment at the rectum and distal portion of the colon, requires sterile preparation, is undesirable for patients, and therefore, leads to compliance issues. Thus, the development of targeted delivery systems for the oral administration of drugs that can reach a specific portion of the intestine is of great interest [79–81]. In this sense, Ling et al. [82] developed protein NPs encapsulated within gastroprotective alginate/chitosan MPs to be administered via the oral route and capable of releasing the NPs in the small intestine and colon in a murine dextran sulfate sodium (DSS)-induced colitis model. NPs were synthesized by desolvation. For this, AvrA (an evolved bacterial protein from Salmonella with anti-inflammatory and anti-apoptotic enzymatic function) was added to the enhanced green fluorescent protein (eGFP) and desolvated with ethanol under constant stirring; eGFP is a protein with no therapeutic function whose loss of fluorescence indicates that changes in the local aqueous environment lead to denaturation. Subsequently, a reducible crosslinker, 3,30 -dithiobis[sulfosuccinimidylpropionate] (DTSSP) (2.2:1 DTSSP:lysine molar ratio) was added to stabilize them and to release the soluble protein in the reducing environment found in the cytosol. Particles were centrifuged and the pellet was resuspended in sterile phosphate buffered saline (PBS) or sterile water. Then NPs, named AvrA NPs, were added to stock 2% w/v alginate solution (dispersed phase). The continuous phase consisted of 1% Span® 80 in mineral oil. Both phases were connected to a microfluidic device, where the collection bath contained 0.5% w/v chitosan and 0.1% w/v calcium chloride dissolved in water of pH 5.5, under constant stirring. Alginate/chitosan MPs were recovered by centrifugation, washed and stored in deionized water until further use. Results showed that the AvrA NPs with eGFP had a diameter of 281  52 nm and PDI of 0.327  0.074. Relatively high PDI values can result in varied uptake kinetics, as larger NPs could be endocytosed more slowly. The negative zeta-potential (11.6  0.6 mV) of these particles could promote accumulation in inflamed regions of the gut due to the repulsion by mucin fibers, and favor diffusion across the dynamic polymeric network of the mucus, thereby targeting the inflamed intestinal tissue with large amount of cationic proteins released by eosinophils. In addition, their small size could allow passive accumulation

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in the inflamed tissue, which is characterized by the loss of the epithelial barrier function and subsequent weakening of the tight junctions, loss of water, and increased permeability of the mucosa. To study the stability of the protein, AvrA NPs were incubated in simulated gastric fluid (SGF) and simulated intestinal fluid (SIF) containing digestive enzymes. Results showed loss of fluorescence in SIF (30%) and SGF (100%). Up to pH 3, eGFP fluorescence can be recovered by adjusting the pH back to 7.4. However, below pH 3, the fluorescence cannot be recovered. The harsh conditions of the stomach would be the major limiting step for the administration of these oral formulations. Therefore, NPs were encapsulated within MPs as oral delivery vehicles. Alginate MPs crosslinked with calcium had an average size of 311  41 μm and a zeta-potential of 12.6  1.9 mV. Therefore, chitosan was introduced in the collection bath to act as a crosslinker in addition to the calcium, and retain the NPs due to increased crosslinking rate and density; chitosan and alginate have complementary electrostatic properties, and form a dense interfacial layer that can prevent the NPs from leaking out. Alginate/chitosan MPs had an average size of 335  50 μm and a zeta-potential of +12.8  7.9 mV. At gastric pH, due to the protonation of chitosan amine groups at low pH, cationic chitosan and anionic alginate allow strengthened interpolymeric associations. On the contrary, at intestinal pH (from 6 to 7.5), alginate hydrogels swell and the charge becomes negative, allowing the release of the encapsulated NPs. To measure the retention of fluorescence, eGFP NPs encapsulated within alginate MPs and coated with chitosan were incubated in SGF, and then washed in PBS to neutralize the pH. Results showed that alginate MPs without chitosan did not retain any eGFP NPs fluorescence, likely due to the pH denaturation and not to the eGFP NPs leakage. A 0.5% chitosan coating was the most effective in retaining eGFP NPs fluorescence. eGFP NPs exhibited a fast, burst release from MPs in SIF after SGF incubation, indicating that MPs are able to protect eGFP NPs in SGF and release them in SIF for subsequent absorption. In addition, a model cell line, HeLa cells, was used to assess NPs internalization. Results showed that eGFP NPs released from MPs achieved an uptake efficiency of 65% when compared to cells that internalized fresh eGFP NPs that have not been encapsulated in the MPs. Finally, unprotected eGFP NPs administered to healthy mice by gavage were unable to transit the small intestine to reach the colon and the majority of the uptake was limited to the jejunum, according to confocal microscopy of the duodenum, jejunum and colon that were harvested from mice sacrificed after 4 h. Conversely, eGFP NPs encapsulated within MPs remained stable and reached the colon. Thus, when using this NP-in-MP formulation, AvrA NPs reduced the downstream macroscopic effects of DSS-induced colitis in murine pre/co-treatment model. Furthermore, an observation from the histology images showed that mice receiving daily gavages of AvrA NPs in MPs had a much lower extent of polymorphonuclear leukocyte presence and infiltration than the other treatment groups (empty MPs and AvrA NPs). After DSS was introduced, all experimental groups exhibited a minor increase in weight gain, followed by a steady decrease until the mice were sacrificed. Mice receiving AvrA NPs in MPs had significantly less weight loss than mice receiving

Delivery platforms for oral drug administration

no treatment. Therefore, this is a strategy that could be used to encapsulate other protein therapeutics to the small intestine and/or colon after oral administration [82]. Mudassir et al. [83] produced pH-sensitive polyelectrolyte methyl methacrylate (MMA)/itaconic acid (IA) nanogels oral administration of insulin. The nanogels were lyophilized in the presence of 2% w/v trehalose as cryoprotectant. In vitro release assays showed that 28.71% of insulin was released in SGF, while 96.53% in SIF. The oral administration of insulin-loaded nanogels (dose of 100 IU/kg) in diabetic rats significantly reduced the blood glucose level to 51.10% after 6 h compared to the control groups. These results suggest that insulin was released in the small intestine [83]. There are numerous strategies (passive and active) to target drug-loaded carriers to specific portions of the GI tract by oral administration. These strategies rely on the expression of receptors and membrane transporters. For example, cells of intestinal epithelium such as M cells, enterocytes or L cells can be targeted with these carriers to improve internalization in the intestinal mucosa. The use of lectin-like receptors to modify the carrier surface can be useful because these receptors bind glycoproteins on M cells and enterocytes. In addition, the use of microbial molecules such as protein invasion or long polar fimbria also can be beneficial because M cells bind pathogen patterns through endocytic pathways. Moreover, apical membrane receptors in M cells such as ganglioside GM1, platelet-activating factor, toll-like receptor-4, glycoprotein 2 and a5b1 integrin are promising targets. Ligands for these receptors can be peptides (Arg-Gly-Asp and LeuAsp-Val) or peptidomimetics. Other molecules such as vitamin B12, vitamin B1, biotin and folic acid can bind receptors expressed on the enterocyte apical membrane, and can be explored to enhance the therapeutic effect of drug loaded nanoparticles after oral administration. The FcRn, expressed by enterocytes of the intestinal epithelium, can also be targeted to allow for the enhanced delivery of orally administered drug carrier systems. In addition, receptors of bacteria on enterocytes in the intestinal flora could be targeted using Ctx or F4 fimbriae on the surface of the carrier. Finally, L cell receptors such as G-protein-coupled receptors can be useful targets for increased uptake of drugs administered by the oral route [84].

3. Conclusions The oral route is the most commonly accepted and widely used for drug delivery due to ease of administration, minimal invasiveness and pain, resulting in high levels of compliance and acceptance by patients. However, many drugs are not suitable for administration by this route due to poor solubility, stability, and/or bioavailability. For example, up to 40% of the drugs on the market and 70% of the compounds under development are poorly soluble in water, which limits their absorption and bioavailability. In addition, compounds such as proteins and peptides are rapidly degraded in the GI tract. The encapsulation of these active compounds in different delivery vehicles, such as NPs, can help overcoming most of their limitations, and allows the targeted delivery of the

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drugs to specific organs or tissues, reducing adverse effects. However, one of the greatest challenges limiting the use of these carriers for de oral delivery of therapeutically active biomacromolecules is the hostile environment found along the GI tract, as the highly acidic pH of the stomach and the presence of enzymes. Accordingly, polymers capable of protecting the active compounds against the drastic GI tract environment and releasing them at a specific site have attracted particular attention. In addition, the ability of the nanocarriers to penetrate quickly through mucus layer covering the intestinal epithelium or to pass through the tight junctions between the epithelial cells has been increasingly investigated. Therefore, the control of the physicochemical properties of the developed nanocarriers, such as size, shape, surface charge and drug release profile is extremely important to ensure therapeutic success after oral administration.

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Further reading Shreya AB, Raut SY, Managuli RS, Udupa N, Mutalik S. Active targeting of drugs and bioactive molecules via oral administration by ligand-conjugated lipidic nanocarriers: recent advances. AAPS PharmSciTech 2018;20(1):15.

CHAPTER 8

(Trans)buccal drug delivery Giuseppina Sandri, Marco Ruggeri, Silvia Rossi, Maria Cristina Bonferoni, Barbara Vigani, Franca Ferrari Department of Drug Sciences, University of Pavia, Pavia, Italy

1. Introduction The buccal route has been investigated for decades as an administration site to achieve drug absorption, to reach systemic circulation or to treat local diseases. The interest has dramatically increased in these years because the buccal route has been identified as an alternative to the conventional oral route [1]. Indeed, the buccal route offers several advantages over the conventional drug delivery routes, including the parenteral route. It is characterized by a high blood supply, which ensures systemic bioavailability, avoiding the hepatic first-pass metabolism and the enzymatic drug degradation. Moreover, the highly vascularized tissue allows a fast onset of systemic effects. In this context, hydrophilic high molecular weight drugs, as peptides or proteins, are favorable candidates [2]. Furthermore, it benefits patient compliance, mainly due to the easily accessibility for the administration by patients, and it is suitable for self-medication both for dosage form administration and its removal [2]. For this reason, it is generally considered in the treatment of chronic disorders when prolonged drug release is required or when the conventional routes do not allow an adequate therapy [3]. However, challenges should be faced to have a successful and a suitable formulation. The residence in the buccal environment should be carefully considered: the accidental swallowing of the delivery systems and the continuous dilution of the saliva could cause a low residence time and consequently a low drug bioavailability. Moreover, the contact between buccal mucosa and the drug delivery systems should be intimate, to allow the drug to act locally or to be absorbed [4, 5]. To go beyond these disadvantages, the drug delivery systems should be characterized by specific interaction with the buccal mucosa, to prevent the drug dispersion via the gastrointestinal system, and should be able to control the drug release and to assure a suitable concentration of the drug in the target tissue. Moreover, it should protect the drug from enzymatic degradation due to the presence saliva or microorganisms [2, 3]. In this perspective, the advancements in nanotechnologies provide tools to overcome these limitations. In particular, nanosystems could enhance drug penetration/ permeation into/across buccal mucosa, avoid drug degradation due to drug encapsulation, and localize drugs close to the site of action/absorption [6].

Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00013-2

© 2020 Elsevier Inc. All rights reserved.

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This chapter aims to describe and discuss the overall advantages and the recent advances of nanosystems for buccal application. A brief introduction focused on the structure and function of oral cavity will be presented, with particular attention to the role of the oral epithelium as barrier against drug penetration/penetration, to saliva, and to the enzymatic barrier. Subsequently, the strategies to effectively deliver drugs in the oral cavity will be assessed, considering the different types of nanocarriers proposed in literature, their features and nanoparticle dosage forms suitable for buccal application.

2. Oral cavity: Anatomic and physiologic features The oral mucosa presents a surface area of about 100 cm2 and consists of two anatomical and functional layers: the epithelium, a thick, stratified squamous avascular layer, and the underlying tissue, a slightly vascular layer of mesodermal origin, called lamina propria [7]. The submucosa region, which is found under the lamina propria, presents a loose and fatty tissue that contains the major blood vessels and nerves of the mucosa [7]. Fig. 1 reports the schematic of the structure of the oral mucosa. Three different types of oral mucosa are recognized: the masticatory mucosa (about 25% of the total mucosal surface), the lining mucosa (about 60% of the total mucosal surface) and the specialized mucosa (about 15% of the total mucosal surface) [8]. The oral lining mucosa (500–800 μm in thickness) is in correspondence of the lips, the cheeks, the soft palate, the ventral surface of the tongue, the floor of the oral cavity and the alveolar processes with the exclusion of the gums. It is covered with a non-keratinized stratified squamous epithelium; it is placed on a lamina propria of loose connective tissue with elastic fibers to guarantee the flexibility needed by the normal function of the oral cavity, as speaking and swallowing [2]. The masticatory mucosa (100–200 μm in thickness) covers the regions mostly subjected to abrasion and shear stress, as gums and hard palate, during chewing. It consists of cornified keratinized epithelium and dense fibrous connective tissue [2]. Finally, the specialized mucosa covers the dorsum of the tongue and it is involved in the sense of taste with the papillae; it is characterized by keratinized and nonkeratinized epithelium. Buccal mucosa refers to the lining of the cheek and the upper and lower lips, which represent one-third of the total oral mucosa surface, although this term sometimes wrongly indicates the whole mucosa of the oral cavity [2]. In Fig. 2 the principal structural features of the keratinized and non-keratinized oral epithelium is reported. The buccal epithelium is approximately 40 to 50 cell layers thick, and the thickness is related with the site. Keratinocytes are the prevalent cell population (approximately 90%), while the remaining 10% is represented by melanocytes, Merkel cells, Langerhans cells and immune system cells (mainly lymphocytes and macrophages) [2, 9]. As a function of the depth of the cell layer, keratinocytes show different patterns of maturation [10]. The basal cells are cubic and capable of mitosis, and this maintains the epithelial population constant in number and thickness: in particular, the tissue

(Trans)buccal drug delivery

Mucus layer

Epithelium

Lamina propria

Submucosa

Periosteum Bone

Fig. 1 Schematic of the structure of the oral mucosa.

homeostasis requires cell differentiation followed by migration from the basal layer up to the superficial one, and the continuous desquamation of the superficial cell layer. The Prickle cells (intermediate layer) accumulate lipids and low molecular weight keratins not arranged in bundles. They present an intracellular lipid portion packaged in the so-called membrane coating granules (MCG) or lamellar granules. These are small

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Fig. 2 Principal structural features of keratinized and non-keratinized buccal epithelium.

organelles that migrate from cytoplasm toward cell membranes where they fuse their membranes and extrude their lipid content (mainly neutral lipids—ceramides and acylceramides—and glycolipids) in the extracellular space. In the keratinized epithelia (especially in masticatory mucosa) the intercellular spaces of the superficial layer are filled with short stacks of lipid lamellae, which fuse at the edges to produce multiple broad lipid bilayer sheets, and a cornified surface layer, like in the skin, is present. In non-keratinized epithelia (especially in buccal mucosa) the extracellular space is filled with amorphous material presenting a low degree of three-dimensional organization in the lamellae. Moreover, in this region, the superficial cells retain their nuclei and some cytoplasmic function, and are surrounded by a cross-linked protein envelope and become larger and flatter [10]. The buccal epithelium is lacking in tight junctions, and it is endowed with gap junctions, desmosomes and hemidesmosomes [2, 10], which are loose intercellular connections. The basal membrane is interposed between the oral epithelium and the connective lamina propria, and it is a complex structure consisting of three different layers: glossy, dense and reticular lamina. The lamina propria is in direct contact with the submucosa or the muscular layer, and can be divided into a papillary layer immediately below the epithelium and a deeper reticular layer, which contains nerves and capillaries needed for blood supplying to connective cells and the keratinocytes. In the lamina propria, native and migrating cells are present, and the extracellular matrix comprises collagen fibers, elastic fibers, glycoproteins and polysaccharides with a ramified structure, in different amounts in the various districts of the oral cavity.

(Trans)buccal drug delivery

The basal membrane represents a mechanical support for the epithelium, and constitutes a selective barrier that allows the exchange of nutrients and catabolites between the epithelium and the connective tissue. Integrity and continuity of the basal membrane are essential in ensuring binding and interaction between the oral epithelium and the lamina propria, playing an important role in the control of growth and differentiation of the epithelial cells [11]. The oral cavity is washed by saliva, a viscoelastic fluid secreted by salivary glands in the mouth. It is composed of 99.5% water and electrolytes (Na+, Ca2+, K+, Mg2+, Cl
, HCO3
, HPO42
), mucins, white blood cells, epithelial cells, digestive enzymes (as amylase and lipase), antimicrobial agents such as the secretory immunoglobulin A (SIgA), and lysozymes. The amount of saliva produced in a healthy person per day ranges from 750 to 1000 mL, with an average flux of 0.5/0.7 mL/min. The saliva volume in the mouth is approximately 1 mL and remains constant due to the continuous swallowing [12]. Saliva is a weak buffer with a pH of 5.5–7 and its composition is affected by the degree of stimulation, as smell, taste and food. Mucins, the characteristic components of the mucus, are highly glycosylated glycoproteins with a molecular weight of about 0.5–20 MDa. They consist of about 500 kDa sub-units jointed by disulfide bridges to give large structures that entrap a large amount of water. This confers to the saliva its viscoelastic behavior, acting as a protective layer over the mucosa. In the buccal environment, mucins are negatively charged, and this is due to the ionization of sialic acid and sulfate residues. These negative charges allow mucins to attach to the surface of the epithelial cells, forming a stable gel layer from 70 to 100 μm thickness that covers all the oral cavity, and mechanically protects the epithelium from trauma during normal activities, as eating, swallowing, and speaking, and avoids mouth soreness [12, 13].

3. Buccal mucosa as barrier for drug penetration/permeation Buccal mucosa is a permeability barrier, and this is mostly imposed by the buccal epithelium acting as a protective layer and as a barrier to the entry of foreign material and microorganisms. However, its permeability is 4–4000 times more than that of skin [2]. The administration of drugs through the mucosa of the oral cavity mainly involves the non-keratinized epithelium so that gingival and hard palatal mucosae are not considered to achieve systemic action. These sites could be useful for the local delivery of drugs to treat localized diseases. To achieve a systemic effect, sublingual mucosa and buccal mucosa are considered. Although sublingual mucosa is more permeable and thinner, with high vascularization, it is generally used for the treatment of acute disorders (as angina pectoris), allowing a rapid onset of the drug absorbed. However, the sublingual mucosa is not suitable for prolonged therapies, since its surface is constantly washed by saliva and the shear of the tongue does not allow the dosage form to remain in the application site and maintain the contact with the mucosa. On the contrary, the cheek (buccal) mucosa is

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less permeable and not suitable to obtain a rapid onset of absorption compared to the sublingual mucosa, but its surface is relatively immobile and more permeable that the other region of the oral cavity (except sublingual one), so it is convenient for drug administration intended for both local and systemic effects. For these reasons, buccal mucosa is the choice for the application of controlled release systems that need to adhere to the mucosa for prolonged periods [14]. The drug transport mechanism through the buccal mucosa involves two major routes: transcellular (intracellular) and paracellular (intercellular) pathways [3]. Fig. 3 reports a schematic of penetration/permeation of drug (red stars) or nanoparticles (green particles) through the buccal epithelium via the transcellular (intercellular) or paracellular (intracellular) pathway. The transcellular route occurs by diffusion of the drug across cell membranes or facilitated diffusion via passive transport, active transport and/or by endocytosis or transcytosis (for macromolecules). The transport of molecules (especially lipophilic compounds and small hydrophobic molecules) via the intracellular pathway is a complex phenomenon, strictly related to the physicochemical properties of the drug, such as its molecular weight, oil/water partition coefficient, ionic charge and structural conformation. The paracellular route, in turn, essentially implies the passive diffusion of the drug through the extracellular lipid domain. It is generally recognized that the lipid matrix of the extracellular space, secreted by mast cell granules (MCGs), plays an important role in the barrier function of the paracellular pathway, especially when the compounds are hydrophilic and are characterized by a high molecular weight, such as peptides [2, 3, 10]. In the buccal epithelium, the intercellular spaces are filled with polar lipids in amorphous state, occasionally stacked in the lamellae. Moreover, the lipophilic cell membranes are surrounded by relatively polar intercellular lipids and hydrophilic cell cytoplasm, and this poses a barrier to polar hydrophilic permeants along with the few tight junctions present. Evidences suggest that compounds mainly permeate via the intercellular lipid domain, although highly lipophilic compounds are generally associated with the cellular membrane lipids [15]. The permeability coefficient (Peff) of a substance can be defined as follows: DP (1) h Where P ¼ partition coefficient: it is determined by the ratio between the solubility of the active ingredient in the epithelial layer and in the vehicle; D¼diffusion coefficient of the drug through the mucosa, that depends on the physicochemical properties of the penetrant and the interactions it has with the epithelial layer; h ¼thickness of the mucosal layer. For ionizable substances, the degree of ionization depends on the pKa of the active ingredient and on the pH present on the mucosal surface. Maximum absorption occurs when the drugs are in the non-ionized form. Peff ¼

(Trans)buccal drug delivery

Paracellular pathway

Transcellular pathway

Fig. 3 Schematic of penetration/permeation of drug (red stars) or nanoparticles (green particles) through the buccal epithelium via the transcellular (intercellular) or paracellular (intracellular) pathway.

The permeability across the buccal mucosa is mainly due to a passive diffusion process, described by Fick’s first law [16]: DP Cd (2) h Where J ¼amount of drug that passes through the mucosa per unit of time and area (flow); D¼diffusion coefficient of the drug through the mucosa; P ¼partition coefficient between mucosa and vehicle; h ¼thickness of the mucosal layer; Cd¼ drug concentration in the vehicle. J¼

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The effective permeability coefficient (Peff) values reported in the literature across the buccal mucosa for different molecules range from a lower limit of 2.2 10
9 cm/s for dextran (MW 4000 Da), across rabbit buccal membrane, to an upper limit of 1.5 10
5 cm/s for both benzylamine and amphetamine, across rabbit and dog buccal mucosa, respectively [17, 18]. Another feature to be considered is the enzymatic barrier represented by enzymes localized on the mucosal surface or in the intracellular compartments [4, 8]. However, this enzymatic barrier is less effective than the gastrointestinal one. Aminopeptidase, carboxypeptidase and esterase were found in homogenates of buccal epithelial cells, nevertheless the use of tissue homogenates did not allow to distinguish between the membrane and cytoplasmatic enzymes [4]. Depending on the transport mechanism (transcellular or paracellular), the enzymatic barrier faced by the drug probably involves part of the above-cited enzymes. In this perspective, a high molecular weight hydrophilic drug as a protein should permeate via the paracellular route, being in contact only with the extracellular domain.

4. Strategies to target the buccal mucosa As described in detail, the buccal mucosa is an effective barrier against molecule permeation. Moreover, the enzymatic barrier, the removal mechanisms as the washing action exerted by saliva, and the mechanical stress caused by the normal activity of the oral cavity as swallowing, chewing and speaking, contribute to impair the drug delivery. Furthermore, the retention times on the absorption site and the application area could be shorter and smaller than required, causing an inadequate drug concentration. Therefore, the different strategies to target the buccal mucosa include the use of penetration enhancers, enzyme inhibitors and mucoadhesive materials.

4.1 Penetration enhancers Penetration enhancers generally are chemical substances able to increase the penetration/ permeation of active ingredients through the mucosa. They are able to reversibly modify the barrier characteristics of the mucosa. Penetration enhancers are generally classified according to their structure, the mechanism of action and the type of drugs [1], and they are able to increase drug absorption without damaging the cell membranes and causing toxicity. The more effective penetration enhancers toward buccal mucosa are surfactants, bile salts, fatty acids and co-solvents [2, 3, 19, 20]. Surfactants (as sodium lauryl sulfate) and bile salts (as sodium glycocholate) are able to increase the permeability of the buccal mucosa, thanks to their capability to interact with the intercellular lipid components [16]. The effect on buccal permeability is mainly related to their capability to disorder intracellular lipids. Furthermore, their activity is related to their concentration. For example, concentrations above the critical micellar concentration (CMC) allow them to disorder extracellular lipids by means of solubilization. Moreover,

(Trans)buccal drug delivery

surfactants and bile salts seem to enhance the penetration/permeation of compounds via the paracellular route. Indeed, their effectiveness is also related to the physicochemical properties of the permeant molecule. In particular, the hydrophilic compounds take advantage of the solubilization and the fluidity of the intercellular lipids, which act as a rate-limiting barrier for hydrophilic molecules; whereas the lipidic molecules mainly localize in the lipidic region of cell membrane, so that the extraction of intracellular lipids has a little effect on the permeability [21]. However, at very high concentrations of surfactants and bile salts, both intercellular (paracellular) and intercellular (transcellular) routes are affected, and this could be due to the disrupting effect that these molecules have toward cell membranes [22–25]. Generally, the enhancement of drug penetration/permeation induced by surfactants and bile salts could cause potential irritation, ulcerations or desquamation, thus nonionic or zwitterionic surfactants are preferred, since they are considered less toxic and less sensitive to changes in pH and ionic strength. Fatty acids are able to increase the fluidity of the intercellular lipids reducing their packaging degree and increasing the partitioning, although these mechanisms have been proved only for skin and not for the buccal mucosa. For the buccal mucosal, the studies are rather deficient: fluorescence anisotropy studies have shown that oleic acid interrupts the organization of membrane lipids in the region of the polar head, suggesting the interaction with membrane lipid domains [26]. Fatty acids are the main lipids in biological membranes, representing about 20% of total lipids in the buccal mucosa, where they are present as free acids, ceramides, triglycerides and phospholipids [19]. Their capability to act as penetration enhancers is related to the type of isomer (cis or trans), the ionization state, the chain length and the branching degree. Saturated fatty acids are generally less toxic than unsaturated fatty acids with the same carbon number. A parallel strategy is the use of co-solvents as propylene glycol, polyethylene glycol and ethanol: these are able to improve the transport of drugs by modifying their solubility and by facilitating the partitioning of the drugs from the vehicle into the mucosa.

4.2 Mucoadhesion Mucoadhesion is an attractive strategy for buccal drug delivery that involves the interaction between the pharmaceutical dosage form and either secreted mucus or a mucosal membrane [12, 27, 28]. Mucoadhesive properties allow a better contact of the formulation with the buccal mucosa and longer residence times. In most cases, mucoadhesion is given to a formulation by the use of polymers. Mucoadhesion occurs firstly by the contact between mucins and mucoadhesive systems (contact stage). In this stage, the hydration, wetting and spreading of the drug delivery systems are fundamental to consolidate the mucoadhesive interaction. The mucoadhesive joint involves the interpenetration of the polymer chains into the mucus layer and the occurrence of polymer-mucin bonding (mainly due to weak bonds and van der Waals, and hydrogen bonds or electrostatic interactions) [2, 3]. Mucoadhesive polymers possess a high number of functional groups, as

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hydroxyl groups or unionized carboxylate groups, capable to form weak bonds. Moreover, the molecular weight and chain flexibility are important to improve interpenetration between polymer chains and mucus, to assure a suitable degree of hydration and chain entanglement with mucins, maintaining system cohesion. Ionizable groups could aid mucoadhesion if they are able to determine electrostatic interactions in the buccal environment with the anionic moieties of mucins [29]. The duration of the mucoadhesion is largely determined by the fast turnover of the mucus layer [28]. Factors as saliva secretion, food intake, local pH, and compositions of the delivery systems, also strongly affect mucoadhesion. In order to improve permeation through the buccal mucosa, scientific research aims at the development of adhesive drug delivery systems, based on polymers with a mucoadhesive behavior. Several mucoadhesive polymers have been used as absorption enhancers. In general, mucoadhesive polymers could be classified as natural (as gums and hyaluronate), semi-synthetic (as chitosans and cellulose derivatives) or synthetic (as polyacrylates). Among different mucoadhesive polymers, chitosan has gained increasing interest in these last decades due to its excellent biological properties, as biocompatibility and degradability. Chitosan is a polysaccharide, copolymer of glucosamine and N-acetylglucosamine, derived from chitin by means of a deacetylation. The mucoadhesive properties of chitosan were studied by Lehr et al. [30]. The mucoadhesive joint formation mainly involves two mechanisms: adhesion by hydration due to water attraction from mucus gel layer [31–33], and chemical bonds formation, in particular hydrogen bonds and ionic interactions between the positively charged amino groups of the polymer and the negatively charged moieties of the mucus gel layer [30–33]. The interactions of chitosan with mucus and its mucoadhesive properties are affected by both physiological factors and the physicochemical properties of chitosan. The extent of mucin adsorbed by chitosan increases upon increasing of the sialic acid residues [32]. Moreover, the mucoadhesive joints are strongly influenced by environmental pH: at acidic and slightly acidic pH, where chitosan is soluble and possesses the highest charge density, the mucoadhesive properties are enhanced [32], while at pH above its pKa (6.5), chitosan mucoadhesive properties are impaired due to its low solubility.

4.3 Enzyme inhibitors Enzyme inhibitors co-administered in buccal drug delivery systems could improve the buccal absorption of drugs, particularly of peptides/proteins. Enzyme inhibitors, such as aprotinin, bestatin, puromycin and some bile salts act by changing the activities of buccal enzymes, altering their conformation and/or impairing enzyme-drug interaction. Furthermore, some mucoadhesive polymers, as polyacrylic acid and chitosans (citrate and EDTA), have been proved to inhibit enzyme activity [34]. In particular, polyacrylic acid (carbomer) is able to bind the essential enzyme cofactors (as Ca2+ and Zn2+): this

(Trans)buccal drug delivery

changes enzyme conformation, causing autolysis and loss of enzyme activity. Moreover, the chitosan salts (e.g., chitosan citrate) or chitosan derivatives (e.g., chitosan covalent bonded with EDTA) allow chitosan to chelate divalent ions and to inhibit metallopeptidases as carboxypeptidase [35]. In recent years, thiomers (polymers as chitosan and polyacrylate derivatives with thiol groups) demonstrated to improve polymer enzyme inhibitory properties [35].

5. The advantages of nanoparticles for buccal delivery Conventional dosage forms are not able to guarantee effective drug therapeutic levels due to the physiological removal mechanisms occurring in the oral cavity, as the washing effect of saliva and mechanical stress, resulting in very short exposure times and unpredictable distribution of drug on the site of action. Currently, one of the more remarkable strategies to improve drug penetration/permeation via the buccal route revolves around the use of nanomedicines. Nanoparticles present many advantages compared to conventional dosage forms. They have a homogeneous distribution (high spreadability) onto the mucosal surface, since nanoscale allows to reduce limitations posed by aggregation. Moreover, they are able to promote bioavailability of poorly soluble drugs, enhancing drug dissolution and controlling drug release. Furthermore, nanoparticles are able to protect the encapsulated drug against the physiological environment. In addition, they demonstrated best mucoadhesive performance when compared to the conventional drug delivery systems, since they possess a high surface area with a marked increase of the interface available for interacting with the buccal mucosa. In this context, the presence of functional groups plays an important role in their interaction with the mucus. For this reason, positively charged nanosystems seem to be favored to form a mucoadhesive joint via ionic interactions, since mucus is anionic due to the presence of negatively charged mucins. However, mucins could form a protein corona to coat the nanoparticle surface and could cause their agglomeration within mucus gel layer. Nevertheless, nanoparticles are generally able to diffuse through mucin chains, and the mucus acts as sieve for these types of systems rather than a surface layer they could deposit on. Therefore, nanoparticles should reversibly interact with mucin macromolecules and should be able to escape from the entangled mucin chain network, and subsequently penetrate across the mucus gel layer to directly interact with the superficial cell layer of the buccal epithelium. For these reasons, nanoparticles should be small enough (around 200 nm) to avoid significant steric hindrance. Moreover, nanoparticles should be suitably engineered to overcome the mucus barrier by escaping from it. Indeed, in these last years, the concept of mucus-penetrating particles appeared. These systems have much more chance to reach the mucosal surface to be taken up, avoiding their retention into the mucus, which presents a pore size of about 340 nm. A decrease in the nanoparticle size usually corresponds to a higher diffusion rate/penetration depth into

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the mucus; however, nanoparticles can be physically entrapped into pores having similar size and characterized by tortuous channels and cages formed by mucin fibers [28, 36, 37]. Moreover, in the case of nanosystems, penetration/permeation enhancement mainly depends on the uptake of the system occurring by means of particle interaction with the surface of the cells and the subsequent repackaging of extracellular lipids. Particle size and zeta potential dramatically affect nanosystem internalization. Indeed, it is described in the literature that nanoparticles having anionic surface charge are characterized by the worst penetration/permeation performance, while the cationic charged nanoparticles are characterized by a better behavior. Moreover, regarding the penetration/permeation into/across the buccal epithelium, 200 nm diameter nanoparticles with neutral zeta potential showed faster and deeper penetration compared to nanoparticles having diameters smaller than 50 nm [28]. Fig. 4 reports a schematic of different types of nanoparticles (polymeric nanospheres and nanocapsules, liposomes, micelles, solid lipid nanoparticles and nanostructured lipid carriers) tested for buccal administration.

5.1 Polymeric nanoparticles Polymeric nanoparticles consist of biocompatible and biodegradable polymers with sizes ranging from 10 to 1000 nm. The drug could be dissolved, entrapped, encapsulated or attached to the polymeric matrix. By varying the method of polymeric nanoparticle preparation, it is possible to obtain nanospheres or nanocapsules. The nanospheres are based on homogeneous polymeric matrices where the drug is uniformly dispersed, while nanocapsules are characterized by an inner compartment surrounded by a polymeric membrane. With these systems, it is possible to modulate the bioavailability and drug nanosphere

nanocapsule

liposome

micelle Nanoparticles for buccal delivery NLC

Fig. 4 Schematic of different types of nanoparticles.

SLN

(Trans)buccal drug delivery

distribution, and the fate of the drug in vivo as a consequence of the type of polymeric matrix (hydrophilicity, biodegradation profile) of choice, and the characteristics of the drug loaded (molecular weight, charge, localization in the nanoparticles) [38]. Different types of polymers have been classified as natural or synthetic, and these include chitosan, gelatin, sodium alginate, albumin, polylactides (PLA), polyglycolides (PGA), poly(lactide co-glycolides) (PLGA), polycaprolactone, poly glutamic acid, poly malic acid, poly(N-vinyl pyrrolidone) [38–40]. Recently, polymeric nanoparticles have been widely studied as promising vehicles for the buccal delivery of insulin. Patil et al. developed insulin loaded nanoparticles based on alginic acid containing nicotinamide as permeation enhancer, by means of mild conditions in an aqueous environment. The nanoprecipitation process developed avoided the use of organic solvents (that could negatively affect insulin stability). Nanoparticles were negatively charged due to the carboxyl groups of alginic acid, and had a mean size of 200 nm, with a low dispersity index. Insulin loading capacities of >95% suggested a high association of insulin with the alginic acid. Insulin stability in nanoparticles was confirmed by HPLC and circular dichroism. When sublingually delivered, nanoparticles with nicotinamide exhibited a higher bioavailability than the subcutaneously injected in a diabetic rat model, suggesting the insulin-loaded alginic acid nanoparticles as a promising sublingual delivery system of insulin [41]. Venugopalan et al. prepared a pelleted nanoparticle system for buccal transmucosal delivery of insulin. Insulin was encapsulated in polyacrylamide nanoparticles prepared by emulsion internal phase evaporation method. Poly(acrylamide) and insulin were dissolved in pH 3.0 methanol, and this polar phase was emulsified by sonication with light paraffin oil containing Span 80. The so formed water-in-oil (w/o) emulsion was stirred continuously to allow the complete evaporation of the internal phase to obtain nanoparticles. The electrophoretic analysis on polyacrylamide gel of the insulin released from the nanoparticles showed insulin structural integrity. In vitro insulin release was related to the polymer concentration used in preparation of the system: the drug release rate decreased by increasing the polymer content. In vivo studies were carried out in rats showing that insulin reached systemic circulation at a controlled rate, without risk of hypoglycaemia [42]. Sandri et al. evaluated the permeation enhancer properties of chitosan nanoparticles in comparison to a chitosan solution. Chitosan nanoparticles were prepared using mild conditions by ionotropic gelation with pentasodium tripolyposphate. The permeability of fluorescein isothiocyanate dextran (MW 4400 Da), used as model molecule for insulin, was carried out by means of Franz diffusion cells, using pig buccal mucosa as biological substrate. The tissue morphology and histology were evaluated after contact with the polymeric systems, using light and transmission electron microscopy (TEM), while confocal laser scanning microscopy (CLSM) was used to localize fluorescein isothiocyanate dextran in the tissue. It was observed that chitosan was able to interfere with the lipid organization in the buccal epithelium, favoring an improvement in the penetration of fluorescein isothiocyanate dextran. Moreover, the nanoparticulate system showed the

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highest apparent permeability coefficient (Papp) value when compared to that of the polymer solution. This demonstrated that, independently of the material considered, the nanoparticulate form has improved properties [43–45]. Suh et al. prepared mucoadhesive nanoparticles, based on chitosan and dextran sulfate, using the ionic gelation method. The nanoparticle diameter ranged from 110 to 360 nm with 0.26–0.44 polydispersion index, indicating that the nanoparticles were homogeneously dispersed. In vitro mucoadhesive properties were carried out by mixing the nanoparticles with a mucin solution in 1:1 weight ratio, and the amount of free mucin was measured using the Bradford protein assay. Ex vivo mucoadhesive properties were evaluated using buccal cells from 10 healthy male and female volunteers. The results obtained showed that nanoparticles were able to interact with mucin and buccal cells, indicating that chitosan- and dextran sulfate sodium salt-based nanoparticles could remain in contact with the buccal mucosa for a long period of time, allowing for drug absorption and an improved therapeutic effect [46].

5.2 Polymeric micelles Polymeric micelles are spherical particles formed by the self-aggregation of copolymeric amphiphiles above their CMC. CMC indicates the concentration value of polymer at which a certain number of monomers aggregate to form micelles. At low polymer concentrations, under their CMC, there are not enough polymer chains to self-assemble into nano-scaled structures, and the polymer acts as surfactant. The most commonly used copolymers are amphiphilic di-block (hydrophilic-hydrophobic) or tri-block (hydrophilic-hydrophobic-hydrophilic) polymers [47]. The amphiphilic copolymers, as polyethylene glycol (PEG), poly(ethylene oxide) (PEO), polycaprolactone (PCL), selfassemble into micelles in aqueous solutions, where the hydrophobic and the hydrophilic portions of the polymer form the core and the shell of the micelle, respectively [48]. The hydrophobic core becomes the reservoir for the lipophilic drug, while the hydrophilic shell affects the micelle stabilization and the interactions with plasma proteins and cell membranes [49]. Polymeric micelles are characterized by higher stability upon dilution than micelles based on surfactants due to their low CMC. Polymeric micelles have been mainly studied for the local treatment of mucosal diseases. In particular, Suksiriworapong et al. evaluated the buccal delivery of itraconazole by thiolated D-ɑ-tocopheryl poly(ethylene glycol) 1000 succinate cysteine (TPGS-Cys) micelles, intended for the local buccal treatment of candidiasis caused by Candida albicans. Micelles were prepared by solvent diffusion and evaporation method. Their particle size and polydispersion index were in the ranges of 8–10 nm and 0.10–0.30, respectively, with a 99% encapsulation efficiency. The release of itraconazole from micelles was biphasic and sustained in simulated saliva solution over 48 h, and micelles showed good mucoadhesive properties toward the porcine buccal

(Trans)buccal drug delivery

mucosa [50]. Other studies described the use of monomethoxy poly(ethylene glycol)poly(epsilon-caprolactone)-graft-polyethylenimine micelles to enhance the solubility and the antifungal effect of Amphotericin B also for the treatment of Candida albicans in the buccal mucosa. The drug and the triblock polymers were co-dissolved in methanol, which was then evaporated to form a thin film. The film was rehydrated in water to obtain micelles. The particle size analysis showed that there were no size differences between the blank and the loaded micelles. Micelles allowed a controlled in vitro release in both normal oral conditions (pH 6.8) and in the presence of Candida albicans infection (pH 5.8) [51].

5.3 Lipid-based nanoparticles Lipid-based nanoparticles are a promising alternative to polymeric nanostructured systems. In fact, the nanoparticles based on polymeric materials present problems related to their production processes, normally involving organic solvents, and difficulties of large-scale production [52]. The use of organic solvents can lead to relative toxicity toward the target cells and high variability of the therapeutic effect. On the contrary, lipid-based nanoparticles are based on highly biocompatible materials and could be manufactured without the use of organic solvents [53, 54]. Moreover, the relative ease of preparation and the low management costs allow a potential development of production on a large scale. Several methods are used to produce lipid nanoparticles, such as high pressure homogenization and ultrasonication, solvent evaporation, solvent emulsification-diffusion, supercritical fluid method, microemulsion method, spray drying, double emulsion, precipitation technique and film-ultrasound dispersion [55]. In particular, high-pressure homogenization is the main technique. In this sense, two different approaches have been set up: high temperature (hot homogenization) and low temperature (cold homogenization). Although the initial step involves the dissolution/dispersion of the drug into the melted lipids, the cold homogenization is much more suitable for temperature-sensitive drugs that could be profitably administered via the buccal route, such as peptides and proteins, as described into detail in the introduction. Lipid-based nanoparticles are conventionally divided into solid lipid nanoparticles (SLNs) and nanostructured lipid carriers (NLCs) [53]. In any case, both lipids and surfactants are usually GRAS (generally recognized as safe) components (completely biocompatible and non-cytotoxic substances approved by Regulatory Agencies) [52]. SLNs consist of solid lipids, such as triglycerides with a high degree of purity, or a mixture of glycerides or waxes. These components give to SLNs a high physical stability, good tolerability, an excellent ability to preserve the active ingredients incorporated against chemical degradation, the possibility to control release kinetics [55], and to incorporate both hydrophilic [56] and lipophilic drugs [57]. However, the crystallization of the lipid

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matrix of SLN could cause active ingredient expulsion during storage. This could be also due to the polymorphic transitions of lipids, generating a contraction of the matrix structure, which tends to expel the loaded active ingredient [52, 56, 58]. To overcome such problems, NLCs have been designed aiming to improve the loading capacity and to avoid the drug expulsion. Three different technology approaches have been developed to optimize the nanostructure of NLCs. The first model, commonly defined as imperfect NLCs, consists of a matrix formed by mixing small quantities of liquid lipids (oils) with solid lipids. The use of lipids with different state at room temperature determines imperfect lipid matrix structures due to the limited packaging of the lipids, offering space for drug molecules [59, 60]. Multiple NLCs are formed when liquid lipids are used at higher concentrations than solid lipids and a phase separation occurs, in which the liquid lipids organize themselves into secondary nanostructures, clearly separated from a continuous phase at solid state. These systems show a high loading capacity, since many drugs are more soluble in liquid lipids. Amorphous NLCs are characterized by a lipid matrix based on lipids having crystallization points higher than the storage and preparation temperatures. Some examples are present in literature mainly focused on the delivery of slightly soluble drugs. Kraisit et al. prepared triamcinolone acetonide-loaded NLCs for buccal local drug delivery. Nanoparticles were prepared by high shear homogenization using spermaceti wax (solid lipid), soybean oil (liquid lipid) and Tween 80 (surfactant) as the lipid phase, whereas water and Tween 80 were used as the aqueous phase. Triamcinolone acetonide-loaded NLCs had particle sizes lower than 200 nm, with a negative zeta potential. The encapsulation efficiency of the NLCs was higher than 80%. In vitro permeation studies showed that triamcinolone acetonide-loaded NLC formulations enhanced drug permeation to a higher extent when compared to the drug in soybean oil [61]. Tetyczka et al. designed domperidone-loaded NLCs based on palmitic and oleic acid as lipid phase, and water containing poloxamer as aqueous phase. The system has been prepared by means of high shear homogenization. NLCs were characterized by 284 nm particle size, with 0.176 PDI and +37 mV as zeta potential. The interaction between saliva and domperidone-loaded NLCs was studied, and no agglomeration phenomena occurred. In vitro permeability on TR146 cells and ex vivo studies showed that NLCs were taken up by buccal cells without cytotoxicity when using concentrations up to 750 μg/mL [62]. Although lipid-based nanoparticles have been proved to enhance drug permeation through the buccal epithelium, they are not generally mucoadhesive, having therefore limited permanence onto the mucosal surface. For this reason, these systems are normally entrapped in mucoadhesive formulations to allow for a longer residence time.

(Trans)buccal drug delivery

5.4 Liposomes Liposomes are spherical vesicular systems, characterized by the presence of phospholipids such as phosphatidylcholine and phosphatidylethanolamine arranged in highly ordered bilayers, having an internal aqueous compartment [63]. They can load either lipophilic drugs into the phospholipid bilayer, or hydrophilic drugs in the aqueous core [64, 65]. Cholesterol is an important component in the liposome preparation: it influences the critical packaging parameter, allowing to obtain a small structure with high stability. Moreover, cholesterol reduces the expulsion of water-soluble molecules caused by diffusion from the bilayer, improving system stability, especially in biological fluids [66]. One of the major advantages of liposomes is their safety profiles: they are based on natural phospholipids and sterols, and possess a composition similar to that of cell membranes [67]. Moreover, liposomes could be deformable, and this could improve their uptake by quickly changing their shape when exposed to the absorptive epithelium [68]. Liposomes are mainly used to enhance drug bioavailability of high molecular weight molecules and poorly soluble drugs, as hereafter described. Liposomes containing lactoferrin and/or lactoperoxidase were developed for buccal local application to control caries in a rat model. Rats were inoculated with Streptococcus sobrinus and fed with a cariogenic diet. In vivo studies carried out for 35 days proved that liposomes were able to decrease the caries incidence to a higher extent than the sodium fluoride (0.2%), lactoferrin and lactoperoxidase solution [69]. El-Samaligy et al. developed silymarin-loaded liposomes to enhance its buccal bioavailability [70]. Liposomes were prepared by the reverse evaporation technique using different molar ratios between lecithin, cholesterol and surfactants, in order to optimize the release profile, permeation and in vitro absorption through chicken cheek. Liposomes having lecithin to cholesterol 7:4 molar ratio possessed optimal encapsulation efficiency. Tween 20 or Tween 80 ratio over 0.5 M reduced the encapsulation efficiency. Release, permeation and absorption studies showed that liposomes containing lecithin, cholesterol, stearylamine and Tween 20 in a molar ratio of 9:1:1:0.5, respectively, gave an enhancement in the drug absorption and permeation compared to the free drug.

6. Dosage forms for the buccal delivery of nanoparticles Although nanoparticulate systems could be administered in liquid forms as colloidal suspensions by means of spraying devices, the majority of drug delivery systems are semisolid/solid formulations (as tablets or solid inserts, films, gels or viscous solutions) aimed to increase the residence time and avoid drug loss due to swallowing. Fig. 5 reports the schematic of different types of dosage forms for the buccal delivery of nanoparticles.

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solid matrix film

Dosage forms for the buccal delivery of nanoparticles

gel

Fig. 5 Schematic of different types of dosage forms for buccal delivery of nanoparticles.

6.1 Films Mucoadhesive films are the preferred dosage form for buccal mucosal administration due to their flexibility, adjustable size and comfort, given by their thickness. Drug loaded nanoparticles are generally blended in a polymer solution before film preparation. Various mucoadhesive polymers have been considered to obtain films as vehicles to carry drug loaded nanoparticles for buccal delivery. Hydrophilic celluloses (e.g., carboxymethyl cellulose sodium and hydroxypropyl methylcellulose), chitosans (e.g., chitosan, carboxymethyl chitosan), polymethyl methacrylate, and gums are taken into account [6, 71–74]. The main challenge is the nanosystem stability after film drying. However, many different types of nanoparticles have been incorporated. For example, in a study by Abd El Azim et al., vitamin B6-loaded liposomes were embedded into a mucoadhesive film based on sodium carboxymethyl cellulose and hydroxypropyl methylcellulose to improve drug permeation [75]. In another study, phospholipid-bile salts-mixed micelles were loaded with cucurbitacin B to address its water insolubility, toxicity, and gastrointestinal side effects.

(Trans)buccal drug delivery

Carboxymethyl chitosan film was used to promote a significant increase in the oral bioavailability of cucurbitacin B in rabbits. This formulation presented the highest values of maximum plasma concentration (Cmax) and area under the plasma concentrationtime curve (AUC), and the lowest time to reach maximum concentration (tmax) when compared to the free drug suspension. Moreover, the matrix did not interfere with the original structure of the micelles [76]. The use of this film containing nanoparticles resulted in a 2.69-fold increase in bioavailability in rabbits when compared to marketed tablets [76]. Morales et al. developed insulin-coated valine nanoparticles embedded in films based on cationic polymethacrylate derivative in combination with hydroxypropyl methylcellulose. The particle coating with the drug avoided drug denaturation during system preparation under high shear forces, as sonication [77–79]. Jones et al. designed SLN-loaded films for buccal delivery of didanosine, a nucleoside reverse transcriptase inhibitor acting as competitive inhibitor of HIV-1 reverse transcriptase. SLNs were prepared using a hot homogenization and ultrasonication technique. Several lipids were investigated to prepare nanoparticles, and glyceryl tripalmitate was identified as the most suitable to obtain small particle sizes (198 nm) and a low PDI (0.175). Buccal films were prepared by casting using hydroxypropyl cellulose and Eudragit [80]. To further increase the interaction and the consequent penetration/permeation of nanoparticles in the buccal mucosa, mucoadhesive polymers can also be used as coatings. Mazzarino et al. loaded curcumin polycaprolactone nanoparticles in mucoadhesive films based on chitosan. Nanoparticles were characterized by a 98% encapsulation efficiency, 250 nm particle size and uniform distribution into the film. The mucoadhesive films allowed to obtain suitable hydration to prolong the drug release in saliva-simulated solution [81]. Furthermore, electrospun films have been recently designed to control drug permeation/penetration in the buccal cavity. These systems combine the adhesion and the retention in the buccal mucosa with a peculiar morphology and porosity that could enhance their effectiveness [82–85]. Reda et al. demonstrated that ketoprofen-loaded electrospun nanofibres based on Eudragit L and Eudragit S were suitable for the local treatment of oral mucositis via buccal administration [86]. Nanofibres were compared with casted films and showed an increase of drug dissolution due to the amorphous state of ketoprofene (a poorly water-soluble drug) in the nanofibers [86]. On the contrary, casted films were unable to achieve a quantitative drug release, since the drug was in the crystalline state in the film. When nanosystems are intended for a systemic effect, films containing nanoparticles can be coated with a backing layer. Such layer can prevent the back diffusion of the nanoparticles into the oral cavity, and avoid drug loss [5]. However, this is useless when aiming at a local effect.

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6.2 Solid matrices To further improve the permanence of the drug in the oral cavity, solid dosage forms as sponges or tablets have been developed. These could also increase drug stability to a greater extent than films. Generally, colloidal suspensions are transformed into powder, mainly by means of freeze-drying. Cryo-/lyo-protectants (e.g., polyalcohols, mannitol and sorbitol, monosaccharides as threalose and glucose, and aminoacids, as glycine) are usually solubilized in the colloidal suspension to protect nanoparticles during lyophilization. For example, Wang et al. prepared buccal tablets loaded with lyophilized naringenin monomethoxy poly(ethylene glycol)-poly(3-caprolactone) nanoparticles for the treatment of oral inflammatory and ulcerative diseases [87]. The tableting process did not alter the nanoparticles, and nanoparticle-loaded tablets allowed for a faster drug release than that of the free drug loaded tablets, as reference. However, in some cases, the tableting can cause a loss of integrity of the nanosystems. For this reason, nanoparticle-loaded solid matrices could be formed directly by means of lyophilization to obtain oral fast dissolving systems. If a viscosizing agent would be added to the colloidal suspension, sponges could be obtained after freeze-drying. Lipidic nanoparticles can deform under mechanical stress, such as the tableting process, and they can undergo lipid phase transitions upon heating during this process as well. For these reasons, Hazzah et al. incorporated curcumin-loaded SLNs into mucoadhesive sponges based on hydroxypropyl cellulose or polycarbophil. Nanoparticles were prepared by high shear homogenization using gelucire containing curcumin as lipid phase, and water and poloxamer as aqueous phase. Curcumin SLNs showed a 298 nm particle size with a 0.476 PDI, 88% encapsulation efficiency and negative surface charge. The sponges prepared by means of lyophilization preserved the nanoparticle structure [88]. In this study, the presence of a lyoprotectant (mannitol) and a plasticizer (glycerol) were crucial to obtain a flexible and elastic structure. Moreover, in vivo results indicate that the residence time in the oral cavity was up to 15 h with polycarbophil as the mucoadhesive polymer [88]. Human studies showed that curcumin SLN-in-polycarbophil sponges have higher Cmax, tmax, and AUC than the SLN-in-hydroxypropyl cellulose sponges. In accordance, the polycarbophil formulation adhered to the mucosa for a longer time (15 h, compared to 4 h), and presented higher matrix porosity and homogeneous distribution of the SLNs. The decreased porosity of hydroxypropyl cellulose sponges decreased swelling and consequent interaction with the mucin. Furthermore, SLNs remained onto the surface of hydroxypropyl cellulose sponges, impairing matrix adhesion to the oral mucosa and releasing the nanoparticles faster than the polycarbophil sponges. In another work, Le et al. analyzed the critical ratio between nanoparticles (in particular SLNs) and powder for tableting. It was evidenced that the excess amount of the solid lipid (stearic acid) in the solid lipid nanoparticles resulted in a lower

(Trans)buccal drug delivery

disintegration rate, which affected the dissolution rate [89]. A small amount of solid lipid in the formulation led to a better hydration of the tablets, resulting in a significantly higher drug penetration through the mucosa than that of systems containing a greater amount of solid lipid [89, 90].

6.3 Gels Mucoadhesive gels have been developed to achieve local delivery in the oral cavity. These systems are well accepted to the patient, and they could easily spread all over the mucosa without disadvantages related to the normal activity of the buccal cavity as speaking. For example, Karavana et al. developed SLNs containing cyclosporine A to be incorporated into a mucoadhesive gel for the treatment of recurrent aphthous stomatitis. SLNs were used to overcome the low bioavailability of poorly soluble drugs as cyclosporine A. A lipid phase based on glyceryl behenate and cyclosporine A, and an aqueous phase based on poloxamer 188 and Tween 80 were separately heated, and then the two phases were mixed using a homogenizer. The particle size and PDI values of Cyclosporine A-loaded SLNs were about 200 nm and 0.37, respectively, with a 95% encapsulation efficiency. In order to obtain a bioadhesive gel, Carbopol or hydroxypropyl cellulose were solubilized in a SLN colloidal suspension. The Carbopol-based gel containing SLNs showed good mucoadhesive properties in vitro. The in vivo studies in rabbits showed that the gel containing cyclosporine A-loaded SLNs allowed for a statistically significant increased rate of mucosal repair, and approximately 70% of cyclosporine A was located in the mucosa after 24 h of treatment, indicating that the cyclosporine A-loaded SLNs were localized in the buccal mucosa [91]. In another work, Franz-Montan et al. evaluated the efficacy of liposome-encapsulated 2% ropivacaine intended for topical anesthesia and its influence on pulpal response. Liposomes were based on egg phosphatidylcholine, cholesterol and α-tocopherol. Liposomes were 370 nm in size, with 0.12–0.17 PDI, and a 27% encapsulation efficiency. Liposomeencapsulated 2% ropivacaine gels based on Carbopol have been administered to volunteers, and showed less pain response to needle insertion and longer tissue anesthesia than the placebo formulations [92]. Marques et al. developed mucoadhesive gels containing NLCs intended for buccal administration of ibuprofen. The lipid phase (glyceride stearate, medium chain fatty acids and ibuprofen) and the aqueous phase (cetrimide and Tween 80) were heated separately, and then the two phases were mixed and homogenized. NLC dispersions were obtained by both ultrasonication and a high-pressure homogenizer. Different processes of preparation did not affect particle sizes, PDI and zeta potential. NLCs were negatively charged with a 97% encapsulation efficiency and a 2.3% loading capacity. NLCs were incorporated in two different hydrogels based on Carbopol 980 or polycarbophil. Rheological measurements showed that both hydrogels exhibited a non-Newtonian pseudoplastic

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behavior. The gel based on Carbopol 980 was characterized by a higher consistency and better mucoadhesive properties than those of the gel based on polycarbophil. Release studies of hydrogels containing NLCs evidenced the capability of these systems to promote a sustained release of ibuprofen [93].

7. Final remarks Nanosystems can reversibly interact with the mucin and can subsequently penetrate across the mucus to directly interact with the superficial cell layer of the buccal epithelium. This allows for the nanoparticle uptake into the buccal cells, enhancing drug permeation/penetration. Moreover, nanosystems can enhance the bioavailability of poorly soluble drugs by increasing drug dissolution, control the drug release profile, and protect the encapsulated drugs against the physiological environment. Furthermore, nanoparticles can be administered as colloidal suspensions or incorporated in a semisolid (gel) or solid (matrix/film) system to increase the retention time in the oral cavity and to obtain an intimate contact with the buccal mucosa. These strategies seem to overtake the main obstacles to buccal delivery, such as limited absorption area and its barrier properties, especially when aiming at a systemic effect, as well as the drug loss due to saliva and normal function of the oral cavity, such as eating, swallowing and speaking. In spite of the great efforts dedicated to this field, the buccal route remains challenging and nanoparticles for buccal administration have not reached the market yet.

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CHAPTER 9

Spray-drying for the formulation of oral drug delivery systems
lder A. Santosb,c Mónica P. A. Ferreiraa, João Pedro Martinsa, Jouni Hirvonena, He a

Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland b Full Professor, Drug Research Program, Division of Pharmaceutical Chemistry and Technology, Faculty of Pharmacy, University of Helsinki, Helsinki, Finland c Full Professor, Helsinki Institute of Life Science (HiLIFE), University of Helsinki, Helsinki, Finland

1. Introduction More than one century ago, the growing need for the fabrication of particle entities with personalized physicochemical properties has fostered the development of sophisticated fabrication methods. In line with this, the spray-drying emerged as an attractive platform technology to mitigate challenges in diverse contexts, such as the food, flavor and pharmaceutical industries [1]. The first patent concerning spray-drying has come to light in the early 1870s [2]. Since that time, this technology underwent constant development and evolution, and became a well-characterized unit operation, ideal for the production of solid dispersions [3]. Its potential in the pharmaceutical field quickly made possible to use this technology in processes of formulation development, particularly for solubility and bioavailability enhancement, improvement of inhaled medicines, stabilization of biopharmaceuticals and tailored release of therapeutic entities [4]. Spray-drying consists in the process of converting an emulsion, suspension, dispersion, or liquid into fine, dispersible particles by the use of an atomizer into a hot drying gas medium, most commonly air [5, 6]. Contrarily to other drying methods, such as freeze-drying, spray-drying is a single-step process, which does not require freezing or high vacuum and, therefore, consumes less energy [5]. Indeed, this powerful technological tool has shown a myriad of advantages over conventional formulation methods throughout the years. Currently, spray-drying is seen as a reliable, reproducible and direct method for particle design and tailored production. Therefore, spray-drying has been extensively used in the fabrication of formulations for the delivery of therapeutics via the pulmonary [7, 8], mucosal (e.g., ocular and nasal) [9–11], skin [12] and oral [13, 14] routes of administration. Among the different routes of drug administration, the oral route is by far the preferred one, and classically the first goal for any newly developed active molecule [15]. Orally deliverable therapeutic compounds overcome a set of drawbacks associated Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00007-7

© 2020 Elsevier Inc. All rights reserved.

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to the use of, for example, uncomfortable and painful injections, the high production costs or the need for strict sterility conditions [16]. Additionally, they are painless and have easy-to-follow therapeutic regimens, having therefore the highest patient compliance [16]. However, the success of oral drug delivery depends on the rate and extent of drug absorption after administration. Here, the dissolution, permeability and solubility of the administered compound play a vital role [17]. According to the Biopharmaceutics Classification System (BCS), these parameters can be characterized by three dimensionless numbers: absorption number (An), dissolution number (Dn), and dose number (D0). These numbers encompass both physicochemical and physiological parameters, and result in the classification of the drugs into four major categories [18]. Interestingly, these parameters have been adopted for setting bioavailability/bioequivalence (BA/BE) standards for the approval of immediate-release oral drug products by the Food and Drug Administration (FDA), the World Health Organization (WHO) and the European Medicines Agency (EMA) [18]. Over the years, researchers have dedicated significant efforts towards the development of nanotechnology-based formulations for oral drug delivery, capable of improving the therapeutic efficacy and reducing the side effects associated with already marketed drugs or biologically active molecules previously considered as undeliverable/undevelopable [19]. Micro/nano-sized drug particles consistently show higher chemical and physical stability compared to other formulations and, most importantly, better solubility and bioavailability [20]. In this sense, spray-drying comes into play as a promising technique for the production of both pure drug particles and drug-loaded micro/ nano-particles made of, for example, polymeric matrices [6]. In this chapter, we review the principle of spray-drying and the different equipment configurations used, with particular emphasis to the main advantages and challenges associated with this technique. Then, we describe the parameters and variables affecting particle formulation, as well as the process parameters to acknowledge for a successful scale-up for industrial production. Lastly, the formulation of drugs/drug delivery systems for oral administration using the spray-drying technology are reviewed. In this context, we provide examples of micro- and nano-particles, as well as pure drug particles that have demonstrated the potential of spray-drying to overcome the limited oral absorption and erratic bioavailability of drug formulations produced by other fabrication methods.

2. Spray-drying 2.1 Principle, equipment configurations, main advantages and challenges Spray-drying is a widely used one-step continuous particle processing operation technique for the conversion of a liquid solution, emulsion, paste or suspension into a solid dry powder, granule or agglomerate. It was described for the first time more than 140 years ago, in a US patent dating 1872, for the improvement of drying and

Spray-drying for the formulation of oral drug delivery systems

concentrating liquids [2]. With a remarkable interest and wide use in the pharmaceutical research and industry, the main principle of spray-drying can be divided into four stages and relies on the: (i) atomization of the liquid feed into very small droplets, which is then (ii) injected into a drying chamber containing hot air (or an inert gas as nitrogen); this causes (iii) instant dry of the atomized droplets into solid particles of specific size that are then (iv) collected to a drying chamber in a matter of a few seconds (Fig. 1A and B) [21–23]. A peristaltic pump is utilized to feed the fluid into the drying chamber at a controlled speed. This fluid passes through an atomizer or nozzle

Fig. 1 Schematic illustration of structured particle production using the spray-drying method and its apparatus: (A) mechanism; (B) general representation of a spray-drying apparatus, (C) co-current flow, and (D) counter current flow configurations of the drying gas in respect to the liquid atomization. The full lines represent liquid atomization and the dashed lines represent drying gas. Panel (A) was created using Servier Medical Art (Creative Commons—Attribution 3.0 Unported—CC BY 3.0). Panels (B)–(D) were reprinted with permission from Sosnik A, Seremeta KP. Advantages and challenges of the spray-drying technology for the production of pure drug particles and drug-loaded polymeric carriers. Adv Colloid Interf Sci 2015;223:40–54.

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(there are different types, such as the rotary atomizer, a pressure nozzle, a two-fluid nozzle or an ultrasonic nebulizer, reviewed in detail elsewhere [21, 23]) and the atomization of the fluid into droplets happens by centrifugal, pressure, electrostatic energy or ultrasound. The size of the droplets is controlled by the type of atomizer being used, also depending on the properties of the fluid of choice (surface tension, viscosity and density). The ultrasonic nebulizer seems to be the most effective in the production of fine and monodisperse droplets. The resulting droplets are subjected to a fast solvent evaporation by traversing the drying chamber with the carrier gas, leading to the formation of dry particles (micrometer scale). Variables in the drying step, such as different local air temperature and humidity conditions inside the drying chamber as a result of non-laminar (i.e., turbulent) gas flow through the system, may lead to differences in the drying degree, and thus, potential non-homogeneity in particle size and morphology. The operation configurations can be open-loop (using a non-recirculating drying gas, usually atmospheric air), being more cost-effective and stable; or closed-loop (using an inert gas, which is recycled and re-used in the drying chamber during the drying process), used to prevent the mixing of explosive gases and drying of oxygen-sensitive substances [6, 24]. The configuration of the spray dryer apparatus chosen will depend on the material properties, dictating the direction of the drying gas flow with respect to the direction of the liquid atomization. The most common configurations are cocurrent and counter current flow, with varying configurations in between [23, 25]. The co-current flow (liquid atomization and drying gas in the same direction, Fig. 1C) allows contact of the final product with cooler air, being useful for drying heat-sensitive materials. In the counter-current flow (liquid atomization and drying gas in the opposite direction, Fig. 1D), the dry product is in contact with the hottest air, and it is desirable when higher thermal efficiency is intended. The dry particles are separated from the drying gas by different collection methods: the bag filter, electrostatic precipitator and the cyclone, being the latter one the most commonly used. The separation done by the cyclone is based on centrifugal forces that arise from the highly rotated air stream, forcing the particles towards the walls of the chamber and into a particle glass collector situated in the bottom of the device. The cyclone is not able to efficiently collect particles smaller than 2 μm, while electrostatic precipitators and filter bag collectors are able to collect particles with sizes down to 100–50 nm [6, 26–28]. Solid products obtained after this process have the advantage of having higher chemical and physical stability compared to liquid formulations, and can be used for the formulation of other pharmaceutical forms, such as tablets or capsules. The spray-drying method presents several advantages [6, 21–24, 29–32], such as: – reproducibility (consistent powder quality throughout the entire process); – single-step, tailorable and controllable processing; – rapid drying of the fluid material into solid droplets;

Spray-drying for the formulation of oral drug delivery systems

– bench-to-bedside translation: reliable scale-up method upon proper process optimization; – different dryer designs for particular applications; – adaptable to heat-labile compounds due to fast drying (seconds or milliseconds) and relatively short exposure time to heat; – prolonged life-span of the obtained dry powder; – versatile: suitable for various types of solutions, suspensions, emulsions, pastes, or melts; – suitable for formulation of poorly water soluble compounds into amorphous solid dispersions; – useful in the formulation and drying of a broad spectrum of compounds, including sensitive compounds and compounds that require controlled protection and release (e.g., ascorbic acid), due to the atomization of the liquid into high-surface area-tovolume ratio small droplets that leads to very fast solvent evaporation; – possibility to prepare and process large biomolecules in a time- and cost-saving manner (compared to lyophilization), although the application of spray-drying for biopharmaceutics still requires extensive optimization and is often coupled to computational and mathematical modeling techniques for a more risk-free process development stage [33]; – powders obtained by spray-drying have better flow properties than conventional formulations; and, – alternative to other drying methods such as freeze-drying, due to shorter time required and cheaper processing, as spray-drying does not involve a deep cooling step [34]. Despite the numerous advantages, the spray-drying technique also presents some challenges [6, 23, 24, 35, 36], such as: – the yield depends on the work scale: at a laboratory scale, with conventional spray dryers, the yield may vary between 20% and 70%; this may result from the loss of particles that get stuck in the walls of the drying chamber, and the low capacity of the cyclone to separate fine particles (> D2

Side view

Template removal

Face view

Doughnut particle

Encapsulated particle

Composite particle

Hairy particle

Porous particle

NP

Fig. 2 Various types of particle morphologies, using the spray-drying method. Adapted and reprinted with permission from Nandiyanto ABD, Okuyama K. Progress in developing spray-drying methods for the production of controlled morphology particles: from the nanometer to submicrometer size ranges. Adv Powder Technol 2011;22:1–19.

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of diffusional motion of the solutes to the radial velocity of the receding droplet surface [40], helping the understanding of the resulting morphology of particles after drying (Eqs. 1–2): Ei ¼

Cs, i eð0:5Pei Þ ¼ Cm, i 3βi

(1)

k 8Di

(2)

Pei ¼

where, Ei is the surface enrichment, Cs,i is the surface concentration of component I, Cm,i is the average concentration of component I, βi is the profile function, Pe is the Peclet number, k is the evaporation rate and Di is the diffusion coefficient of solute i. Vehring et al. [41] showed that while low Pe (1) indicates that the recession of the surface is fast compared to the diffusional motion of the dissolved molecules, leading to a fast increase in concentration of the solute at the surface as the evaporation progresses. This causes a local increase in viscosity, with subsequent surface skin or shell formation that, being thick enough, can maintain the shape of the particle. The remaining solvent diffuses through the shell or evaporates through openings in the shell. Ultimately, this may lead to a hollow particle that can collapse or wrinkle, depending on the thickness and mechanical properties of the skin. The Pe varies depending on the material properties of solutes and solvents, as well as with process factors that influence the evaporation rate. The Pe is not constant over the drying process as the diffusion coefficient and evaporation rate are not constant, especially when there are two or more solvents involved, or there is a phase transition of solutes (e.g., crystallization). Thus, understanding the process and formulation factors that influence the Pe allows the achievement of the desired particle properties [40]. 2.2.2 Feed solution properties When formulating a certain active pharmaceutical ingredient (API) into particles by spray-drying, several factors need to be taken into account regarding the choice of the excipients, so that particles with the desired properties can be obtained. Excipients like polymers, sugars or even micro/nano-particles are required when dissolution rate of the API needs to be tailored. Thus, solubility and chemical stability of the API and excipients in a certain solvent(s) are of great importance. Moreover, there are other properties to consider when selecting the feed solution constituents [43, 44]. Glass transition temperature (Tg) is important for the selection of the spray-drying temperature, for the interactions between API and excipients, and when aiming at the API in an amorphous form. This is because at room temperature, excipients with high Tg possess a glassy state with

Spray-drying for the formulation of oral drug delivery systems

great viscosity, which reduces API mobility and thus, risk of recrystallization [45]. The nature of interactions (hydrogen bonds, electrostatic, ionic or hydrophobic) and the presence of functional groups in both API and excipients is critical for the development of dry particles with the desired properties. Strong interactions between the API and the excipient provide higher miscibility between them [46]. Hygroscopicity may affect the stability of the formulation during storage, especially in the case of amorphous systems. Low miscibility of the API and the polymer, as well as decreased Tg of the formulation affected by presence of moist in the fluid, are potential mechanisms that change the viscosity of the system, leading to drug-rich phases that endorse phase separation and recrystallization [47, 48]. Molecular weight and thermal stability are also factors to take into account for particle formulation in spray-drying [21]. Solvent properties are equally important in the selection of the appropriate solvent for the spray-drying and formulation process. The most important criteria for solvent selection rely on: [45] – high solubility of drug and excipients in the selected solvent; – final solution viscosity, crucial for an effective atomization step; – low toxicity and environmental hazards; – high volatility for fast evaporation during the drying step, since the evaporation rate determines the properties of particles with respect to the morphology and radial distribution of the components; – chemical stability of the excipients and API in the solvent; and – non-combustion in spray-drying environments. 2.2.3 Process parameters Process parameters are determinant in obtaining the desired particle formulation. The feed concentration determines the particle fraction that becomes solid upon evaporation of the solvent. Higher concentrations of feed lead to a lower amount of solvent per droplet, a faster evaporation, the Pe values become higher, and the final product results in low density porous particles and rough surfaces [49]. The feed rate controls the feed entering the drying chamber, and can alter the evaporation rate, particle size, morphology and density. High feed rates are associated with higher moisture contents in the final product and with larger particles, if the atomizing gas flow is not adjusted accordingly. Higher yields are reported when a lower feed rate is set [26, 50]. The inlet temperature has a direct effect on the heat and mass transfer during the drying of the droplets, as the evaporation of solvents will occur faster or slower. A higher inlet temperature leads to altered morphologies, surface roughness, porous or collapsed particles [40, 41, 51], while lower inlet temperatures slow the evaporation rate and may result in stickier particles, which may lead to a lower yield due to the stickiness of the powder to the cyclone walls [51]. The morphology and size of particles are directly influenced by the atomizer used in the spray-drying equipment, as well as

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by the type and flow rate of the drying gas. Depending on the type of atomizer utilized (see section 2.1), the droplet mass median diameters range from less than 10 μm (important for pulmonary and parenteral administration routes) to 100 μm, which translates into a typical dry particle diameter range of 0.5 to 50 μm [6]. The properties of the final product are imparted by density and specific heat capacity of the drying gas utilized. N2 gas with lower density (1.1233 kg/m3) has been shown to produce smaller particles with different surface morphologies and to affect more the crystallinity of solutes compared to CO2 (1.7730 kg/m3), as CO2 demonstrated higher efficiency in the drying process [51–53].

2.3 Scale-up considerations Spray-drying has progressed towards a mature technique for industrial scale, allowing the production of tons of particle formulations per day, although the production of high-value protein powders still remains a challenge and is limited to the laboratory scale. However, one of the biggest barriers to the translation of a spray-dried product to the market is the scale-up. Often, the biggest challenge is resumed to obtaining the same particle size and structure as in the laboratory scale, due to the significant interaction between process conditions, product structure and properties. Scale-up in spray-drying is made difficult with the high costs of industrial-scale equipment, the need of new optimizations to tune the desired particle size distribution and morphology, to keep reproducibility from the laboratory scale to the industrial scale during the scale-up process, and the particle recovery/yield may be limited to the type of collector [40, 54]. This led to the development and utilization of fundamental computer modeling-based particle design approaches to minimize trial-and-error, as well as to save time and costs, helping to improve control over the product properties [4, 33, 55]. Practical experience and knowledge of the operator regarding process and component scale-up, properties of the product from the pilot-scale experiments and effects of variables like heat-mass transfer and humidity are also very important for the successful scale-up to industrial scale [4]. In spray-drying scale-up, it is important to acknowledge the process parameters used during the small-scale particle development and production, in order to achieve the desired particle properties in the larger scale: – feed properties, concentration and feed moisture content can be adjusted at the larger scale to match drying time; – atomized droplet size distribution is a very important and critical feature that needs to be taken into account, as it controls droplet size and hence particle size; in addition, small scale spray dryers usually come with two-fluid nozzles that produce small particles, whereas large scale spray dryers use pressure nozzles for obtaining larger particles [56];

Spray-drying for the formulation of oral drug delivery systems

– the droplet drying trend distribution and residence times in the drying chamber tend to be longer for bigger scale spray dryers; extra drying time is often required, which has an impact in the product’s particle size, morphology, and residual solvent/moisture content; – the desired particle/droplet collision history and the formation or not of agglomerates is also important to avoid wall contacts and build-up at all scales; and – yield and heat loss from the instrument walls are variable from laboratory to industry scale: due to a large fraction of fine particles lost in the filters and formation of wall deposits, yields obtained with the small-scale spray dryer are up to 60–70%, largely because of incomplete drying and poor thermal efficiency; large-scale dryers have better thermal efficiency, as deposits are decreased by adjusting the operation to attain hygienic or even aseptic conditions [57]. There are different commercially available spray dryers for laboratory scale, with the Nano Spray Dryer B-90 (B€ uchi, Labotechnik AG, Switzerland) currently being one of the most used state-of-the-art spray dryers. Such device may achieve particle sizes ranging from 300 nm to 5 μm, with satisfactory yields, and adequate particle sizes for drug delivery applications in the oral, intravenous, transdermal and pulmonary fields [28]. The Mini Spray Dryer B-290 however, enables the more straightforward scale-up of the process, initially to pilot and later onto industrial production, due to its larger capacity of production, although the particle size collection is limited to the micrometer range (1–25 μm). Spray dryer equipment and scale-up process considerations are thoroughly described elsewhere [4, 6, 23, 28, 56, 58]. Some pharmaceuticals produced by spray-drying are available in the market for the treatment of diseases like high cholesterol [59], cystic fibrosis [60], hepatitis C [61] and tuberous sclerosis benign tumors [62] as amorphous solid dispersions, often formulated as tablets or granules for oral intake [21].

3. Formulation of drugs/drug delivery systems for oral administration using the spray-drying technology The formulation of drugs into particulate systems (micro- and nano-particles) presents itself as a promising alternative to overcome biopharmaceutical problems, such as poor water solubility, stability issues, transport barriers, as well as to target and tune the release of drugs [63]. As an example, formulating poorly water-soluble molecules into particle entities increases the drug surface area available for interaction with solvent molecules, improving its dissolution rate compared to the raw material, absorption, and thus, the bioavailability in the blood stream, avoiding the administration of high drug doses [64, 65]. Spray-drying has been applied in oral drug delivery mainly to improve the solubility and controlled release kinetics of therapeutics of interest, by decreasing the size of the drug particles and increasing their surface area-to-volume ratio available for contact

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with the solvent. Spray-drying has also been used to improve drug wettability induced by the use of a hydrophilic polymer, and/or changing the crystalline structure of the drug into a more favorable amorphous state, easily achievable by formulating them with excipients of interest into micro- and nano-particles [63, 66, 67]. In the next sections, we describe the most commonly used excipients to improve the properties of drugs prepared by spray-drying to be delivered orally, and the importance of the spray-drying technique compared to other methods of particle preparation. Additionally, the importance of micronization by spray-drying technique is highlighted, with examples of drug micro- and nano- particles, and pure drug particles for application in oral drug delivery (Table 2).

3.1 Conventional methods versus spray-drying of micro- and nano-particle formulations Conventional micro- and nano-particle production methods and technologies usually vary from top-down (homogenizers [84] and wet-milling [85]) to bottom-up approaches (confined liquid impinging jets [86] and high-gravity-controlled-precipitation reactors [87]), with the aim of obtaining small, stable and monodisperse particles, features that are required for pharmaceutical formulations and most biomedical applications. Conventional top-down approaches bring together a series of issues, including high risk of milling media contamination, and there is a number of critical process parameters that influence the profile of the final micronized drug particles and nanosuspensions [85]. Bottom-up approaches usually involve the use of solvent-anti-solvent combinations, sometimes requiring the use of hazardous organic solvents in large volumes, and may lead to particle aggregation [88]. Also importantly, top-down and bottom-up methods of particle production often yield micro- and nano-particles in the form of suspensions, requiring an additional drying step to obtain the final particle powder (usually by spray or freezedrying). In this scenario, the spray-drying technique has become one of the most powerful tools to obtain dry particles with the desired properties. The effortful investments and constant development of this technique made it extremely appealing in both laboratorial and industrial setups for the production of dry powders over a number of conventional methods. Hence, and taking into account the advantages described in Section 2.1, the spray-drying technology seems to be a game-changer. Indeed, it allows for the production of particles with controlled size, polydispersity and morphology, and avoids aggregation of the final products, due to the combination of both particle engineering/formulation and fast drying in one single step [6]. Additionally, it is reproducible, scalable without major modifications, and provides a prolonged lifespan of the final products due to the efficient dehydration occurring at the drying phase [6, 89].

Table 2 Examples of micro-/nano-particles and pure drug particles produced with spray-drying for oral drug delivery applications. Spray dryer of choice

Microparticles

Application

Size

Main conclusions

Refs.

B€ uchi Spray Dryer B-290

Insulin encapsulated-alginate microparticles for assessment of a developed ELISA AKT assay for in vitro evaluation of insulin structure and activity

2.1 μm

[68]

B€ uchi Mini Spray Dryer B-290

Bioadhesive chitosan-based microparticles encapsulating metformin hydrochloride for antidiabetic therapy

2–33 μm

Nano Spray Dryer B-90

Mucoadhesive microspheres of sitagliptin for antidiabetic therapy

2–8 μm

B€ uchi Mini Spray dryer B-191

Cell lysate of 4T07 murine breast cancer cells incorporated in an HPMCAS and EC polymer matrix for oral breast cancer vaccine immunotherapy

1–4 μm

B€ uchi Mini Spray dryer B-191

Inactivated influenza A/PR/34/8 H1N1 virus incorporated in a Eudragit S and trehalose polymer matrix for oral vaccination against influenza virus

1–6 μm

Spherical or varying dimpled particle morphologies; Bioactivity of insulin was preserved, with minimal impact on its structure during the spray-drying process Bioadhesive properties of the spraydried microparticles: High encapsulation efficiency of metformin, low moisture content and homogenous particle size distribution; Increase in metformin retention on porcine mucosa with increasing chitosan:metformin ratios High drug loading and percentage yield (73 0.2% and 92 0.3%, respectively); Excellent mucoadhesion (7.8 0.3 h); Sitagliptin was retained in the GIT for an extended period of time (12 h) compared to control (4 h) in vivo Enteric protection and sustained-release profile of antigens; Higher numbers of CD4+ cells and serum antibody titers in vaccinated animals; Vaccinated animals developed significantly smaller tumors Yields of about 75%, antigen release over 50% after 5 h; In vitro hemagglutination activity assay promoted antigen stability antigen with no loss of HA influenza protein activity; Microparticle oral vaccination in mice resulted in enhanced antigen-specific IgG antibodies; Increased levels of protection in vaccinated animals upon challenge with homologous and heterologous influenza viruses

[69]

[70]

[13]

[71]

Continued

Table 2 Examples of micro-/nano-particles and pure drug particles produced with spray-drying for oral drug delivery applications—cont’d Spray dryer of choice

Application

Size

Main conclusions

Refs.

B€ uchi Mini Spray dryer B-191

Microparticle-encapsulated inactivated Vibrio cholerae, in an Eudragit—L30D-55 or FS30D polymer matrix

3 μm

[72]

B€ uchi Mini Spray dryer B-191

Plasmid DNA to hepatitis-B surface antigen incorporated in albumin-based chitosan microparticles as oral vaccine against hepatitis B

0.8–1.5 μm

B€ uchi 290 Mini Spray Dryer and B€ uchi 190 Mini Spray Dryer

Two times spray-drying of encapsulated budesonide into a crosslinked polyelectrolyte chitosan-alginate polymeric microparticle coated with Eudragit for colitis therapy

4.045 μm

B€ uchi 190 Mini Spray Dryer nozzle-type

Ibuprofen incorporation in gelatin microcapsules for improvement of oral bioavailability of ibuprofen.

6.34 μm

Eudragit L30D-55 microparticles released 86% after 24 h, whereas FS30D released less than 30% of Vibrio cholerae in neutral medium; Antigenicity and microparticle morphology affected by the inlet temperature; Rats inoculated with microparticles exhibited vibriocidal antibody titres Albumin microparticles provided protection of DNA from nuclease degradation; Oral immunization of BALB/C mice increased the titer level of both IgA and IgG versus subcutaneous immunization Limited budesonide release at pH 2 and 6.8, and higher release at the colon (pH 7.4); Colitis severity was suppressed in vivo upon oral administration of budesonide microparticles Enhanced dissolution rate of ibuprofen within the gelatin microcapsule compared to ibuprofen powder at pH 1.2 due to change in the crystallinity towards an amorphous state, improving absorption rate; Significantly higher initial plasma concentration, Cmax and AUC of ibuprofen when encapsulated within the gelatin microcapsule

[73]

[74]

[75]

B€ uchi Mini Spray dryer B-191

Didanosine-loaded PCL microparticles for antiretroviral therapy

36–118 μm

B€ uchi Mini Spray Dryer B-290

ABZ-containing chitosan-pectin-CMC microspheres for improvement of oral absorption and bioavailability for anthelmintic therapy TMS-PVP microparticles for improvement of its solubility, dissolution and oral bioavailability for hypertension therapy

2.8 μm

B€ uchi Mini Spray Dryer B-290

Chitosan microcapsules cross-linked with TPP for oral delivery of venlafaxine (highly soluble drug) for anti-depression therapy

3–10 μm

Nano Spray Dryer B-90

Gelatin nanospheres for the delivery of vildagliptin for antidiabetic therapy

445 nm

B€ uchi Mini Spray Dryer B-190 nozzle-type

Nanoparticles

N.D.

Encapsulation efficiency from 60 to 100%, the yield of up to 65%; Drug released within 120 min in vitro; Drug degradation during gastric transit prevented; Oral administration to rats increased 2.5-fold the oral bioavailability compared to a didanosine aqueous solution ABZ microspheres with optimized and predictable properties; Increased bioavailability of ABZ when administered as ABZ microspheres Crystallinity of TMS in the amorphous state when encapsulated in the PVP microparticles; Improvement in aqueous solubility (40,000-fold) and dissolution rate (3fold) compared to TMS powder Chitosan-TPP ratio affects the release of venlafaxine; Low-viscosity chitosan and a chitosan: TPP ratio of 1:1 imparted a controlled release of venlafaxine (maximum release at 6 h in SIF of 60%) Excellent mucoadhesive properties; No changes in the physicochemical properties; Can be stored below RT; Retained in the GIT to a larger extent both in vitro and in vivo compared to the pure drug

[76]

[77]

[78]

[79]

[80]

Continued

Table 2 Examples of micro-/nano-particles and pure drug particles produced with spray-drying for oral drug delivery applications—cont’d Spray dryer of choice

Application

Size

Main conclusions

Refs.

Nano Spray Dryer B-90

Vildagliptin and metformin hydrochloride-loaded carbopol NPs for antidiabetic therapy

437 nm

[81]

MIVM integrated with a SD-05 Spray Dryer

SR13668-loaded PLGA NPs for cancer prevention

150 nm

B€ uchi Mini Spray Dryer B-290

EudL-coated or EudS-coated BSA nanospheres loaded with cyclodextrin complex for IN delivery (anticancer and anti-inflammatory) Celecoxib/solid ethyl cellulose solid dispersions for the treatment of osteoarthritis and rheumatoid arthritis, and for anticancer therapy

170–400 nm

Stable for up to 12 months at RT, and controlled RH; Stomach-specific drug delivery; Prolonged residence time in the stomach Long-term stability; Fast production (around 10 min); Allows for adjustment of nanoparticle/ drug concentration before administration; Significant increase in the SR13668 oral bioavailability compared to tablets Increased IN solubility; Zero-order drug release kinetics; Negligible release at acidic pH, without limiting drug availability at pH 7.4 Size and morphology maintained intact after spray drying; Improved oral bioavailability; Faster drug absorption compared to commercial capsules Good solidification and no stickiness; Flurbiprofen in a changed amorphous state in the NPs; Drug solubility improved by about 70,000-fold; Enhanced oral bioavailability in vivo Mass recovery above 60%; PCL underwent amorphization during the spray-drying process; Increased plasma concentration and oral bioavailability

Niro Mobile Minor Spray Dryer

100–150 nm

N.D.

Flurbiprofen-loaded NPs for the treatment of rheumatoid arthritis and other rheumatic disorders

300 nm

B€ uchi Mini Spray Dryer B-290

EFV-loaded PCL NPs for the treatment and prevention of HIV/AIDS

207 nm

[14]

[65]

[64]

[67]

[30]

Pure drug particles

B€ uchi Mini Spray Dryer B-190 nozzle-type

RXF-loaded solid dispersion NPs for the prevention and treatment of a wide range of diseases

180 nm

SD-1000 Spray-Drier

Nitrendipine nanocrystals for antihypertensive therapy

175 nm

N.D.

NanoCrystal spray-dried powder of cilostazol for antithrombotic therapy

220 nm

Does not require the use of an organic solvent; Adequate for the reduction of the drug particle size and to change the drug from crystalline to amorphous form; Increased drug solubility and dissolution rate; Increased oral bioavailability in vivo Nitrendipine remained in the crystalline state; 100% of the drug dissolved within the first minute; Increased maximum peak concentration in plasma compared to physical mixture and commercial tablet; Nanocrystals re-dispersed without aggregation or agglomeration; No changes in the particle size after spray-drying; Improved dissolution rate and an increased oral bioavailability Dissolution rate of 95% after 60 min; Increased absorption rates compared to the hammer milled, jet milled techniques; Increased oral bioavailability in vivo

[66]

[82]

[83]

ABZ, albendazole; AIDS, acquired immunodeficiency syndrome; AKT, protein kinase B; AUC, area under the curve; BSA, bovine serum albumin; CD, cluster of differentiation; CMC, carboxymethylcellulose; EC, ethyl cellulose; EFV, efavirenz; ELISA, enzyme-linked immunosorbent assay; EudL, Eudragit L-100; EudS, Eudragit S-100; GIT, gastrointestinal tract; HA, hemagglutinin; HIV, human immunodeficiency virus; HPMCS, hydroxyl propyl methylcellulose acetate succinate; IgA, immunoglobulin A; IgG, immunoglobulin G; IN, indomethacin; MIVM, multi-inlet vortex mixer; N.D., no description; NPs, nanoparticles; PCL, poly(epsilon-caprolactone); PLGA, poly(lactic-co-glycolic acid); PCP, polyvinylpyrrolidone; RFX, raloxifene hydrochloride; RH, room humidity; RT, room temperature; SR13668, [2,10-Dicarbethoxy-6-methoxy-5,7-dihydro-indolo-(2,3-b)carbazole]; TMS, telmisartan; TPP, sodium tripolyphosphate.

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Altogether, these advantages can certainly foster the bench-to-bedside translation of spray-dried products, and therefore, spray-drying has been thoroughly explored for the encapsulation of different cargos, such as pigments, food ingredients, aromatic oils, and specially drugs, within different types of carriers like polymeric nano- and microparticles, micro- and nano-composites, as well as in the formulation of pure drug particles [30, 34, 90, 91].

3.2 Common excipients used in the preparation of drug formulations prepared by spray-drying The use of certain excipients helps the formulation process and allows the shaping of the desired properties of the final product in spray-drying for suitable oral drug delivery. The most commonly used excipients for the preparation of particulate systems by spraydrying for oral drug delivery include polymers, cyclodextrins, counter-ions, oils and surfactants, envisioning the improvement of the solubility and/or therapeutic effect of drugs meant to be delivered orally. Some polymers commonly used in oral drug delivery include hydrophobic polymers, such as poly(lactic-co-glycolic acid) (PLGA) [14] or polycaprolactone (PCL) [30], and hydrophilic polymers like alginate [68], chitosan [69, 79], poly(vinylpyrrolidone) (PVP) [66], and gelatin [75]. Hydrophobic polymers and poorly water-soluble drugs are initially dissolved in organic solvents and further submitted to spray-drying. Using organic solvents is an advantage, as many have lower boiling point than the polymer’s melting temperature (Tm) or Tg. For example, PLGA (Tg ¼ 40–60 °C) and PCL (Tm ¼ 55–60 °C) present higher Tg and Tm than the boiling temperature of acetone (56 °C) or dichloromethane (40 °C), which allows for evaporation at low temperatures, and avoids agglomeration and sticking of particles [89, 92]. On the one hand, using solvent combinations is also convenient to alter the solubility of the polymer and therapeutic cargo being formulated, such as to obtain a boiling point that yields particles with desired properties. On the other hand, using water as a solvent is preferred due to its non-toxic nature, in addition to the lower risk of explosion. Combining water and water-miscible solvents like ethanol is also done to adjust the final boiling point and to enable spray-drying at lower temperatures [14, 66, 93]. Moreover, polymers are used to modulate the release profile of the encapsulated therapeutic cargos, to protect the cargos from harsh environments, such as the stomach pH, while allowing their release in the intestine, where the pH is higher, to increase the interaction between particles and mucosa, as well as protect delicate cargos (proteins and genetic material) (Fig. 3). These strategies can promote an increased and more efficient drug absorption in the desired place, leading to a more stable and possibly prolonged bioavailability of the drug in the blood stream (Fig. 3) [74, 94, 95]. Cyclodextrins are used as excipients to convey a better dissolution rate by entrapping poorly water-soluble drugs within their hydrophobic core, while their hydrophilic

Spray-drying for the formulation of oral drug delivery systems

Fig. 3 Schematic representation of the main advantages of using polymers (e.g., chitosan) as drug carrier systems for oral drug delivery. After being orally administered, polymeric nanoparticles are able to protect the drug from the harsh stomach environment, and allow for its targeted release in the intestine, subsequently favoring drug absorption and leading to an increased bioavailability of the drug in the blood stream. Reprinted with permission from Mohammed AM, Syeda TMJ, Wasan MK, Wasan KE. An overview of chitosan nanoparticles and its application in non-parenteral drug delivery. Pharmaceutics 2017;9.

moieties are exposed to the outside aqueous environment. The inclusion of poorly water-soluble cargos in the core of cyclodextrins was reported to alter the crystalline state of the cargos towards an amorphous state, promoting a higher dissolution rate and enough stability for storage [96]. Counter-ions are employed for the formation of soluble salts of initially poorly watersoluble drugs (or displaying a pH-dependent solubility) and their choice depends on the nature (acidic or basic drug) and the pH-solubility profile of the drug with different counter-ions. Salt formation of poorly water-soluble drugs increases the dissolution rate by enhancing the apparent solubility in the diffusion layer/microenvironment immediately around the spray-dried solid particle [97, 98]. Additionally, emulsifying agents like oils and surfactants, together with the right solvents and co-solvents, are very useful in the formulation of solid self-emulsifying drug delivery systems, commonly used in the formulation of lipophilic drugs.

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Such formulations can be included in oral solid dosage forms, such as tablets or capsules, and return to liquid emulsions upon reconstitution. This exposes the poorly soluble drug to the aqueous environment, prone to a better dissolution and absorption. Reports state a great increase in solubility, stability and improvement of oral bioavailability of poorly water-soluble drugs like non-steroid anti-inflammatory APIs [99, 100] and clopidogrel [101] by formulating them using this strategy.

3.3 Preparation of particulate systems with spray-drying and applications in oral drug delivery 3.3.1 Microparticles Microparticles have been widely developed to overcome certain physicochemical properties of drugs that limit their bioavailability. Since oral drug delivery does not necessarily require nanoscale-sized materials, microparticles come across as very promising drug carriers for oral administration, as they can be easily obtained by spray-drying. The cyclone powder collection system, having a cut-off particle size of minimum 2 μm, continues to be one of the most popular and successful particle collector options, and thus, it is commonly chosen for the production of micrometer-sized particles for oral drug delivery applications. For example, bioadhesive chitosan-based microparticles were prepared by spray-drying, using the B€ uchi Mini Spray Dryer B-290, equipped with a cyclonic particle collector, for the delivery of the model antidiabetic and highly water-soluble metformin at the oromucosal/buccal cavity. Metformin, commonly administered as a tablet, possesses a narrow oral absorption window and leads to known gastrointestinal side effects. In this study, the microparticles presented a wrinkled to spherical morphology and sizes varying from 8 to 16 μm, depending on the ratio of chitosan and metformin used. The authors demonstrated that the bioadhesive chitosan microparticles encapsulating metformin were successfully prepared by spray-drying and increased the retention of the drug at the oromucosal cavity. This microparticle system may extend the drug absorption at the buccal level due to the transmucosal drug delivery effect, potentially improving the drug bioavailability and limiting the gastrointestinal-related side effects of metformin, since less drug is swallowed when using bioadhesive formulations [69]. Also importantly, the system could be utilized for other drugs suitable for oromucosal administration. The preparation of drug-encapsulated microparticles by spray-drying technique has been very useful in the development of novel drug formulations for numerous therapeutic applications via oral administration, such as diabetes [68–70], cancer immunotherapy [13], oral vaccines [71–73], inflammatory bowel diseases [74], antiinflammatory therapy [75], anti-infectious therapy [76, 77], cardiovascular diseases [78] and even depression therapy [79] (Table 2). For example, Chablani et al. [13] have developed a potential prophylactic oral vaccine for cancer immunotherapy targeting breast cancer. They incorporated the whole cell lysate of 4T07 murine breast cancer cells

Spray-drying for the formulation of oral drug delivery systems

as a source of antigens in a hydroxypropyl methylcellulose-ethyl cellulose-β cyclodextrin matrix (water-soluble) and spray-dried to formulate an enterically protected vaccine microparticle to be administered orally, free of organic solvents, and yielding microparticles of 1–4 μm diameter in size. The aim was to deliver these microparticles carrying breast cancer antigens to the M-cells in the small intestine for immunization, also carrying aleuria aurantia lectin for targeting of these cells [102]. The microparticles remained intact in gastric conditions and released 30% of their content in intestinal fluid after 8 h, in a diffusion-controlled release from the polymeric matrix, suggesting that the microparticles remain fairly intact even in the intestine, an important factor for an efficient antigen presentation to the M-cells. Upon several booster doses of these vaccine particles, results showed both humoral and cellular immune response in vaccinated mice. Post-vaccination animals were challenged with live 4T07 cells, developing significantly smaller tumors than control animals. By the use of an easy and scalable fabrication method for microparticle vaccine preparation such as spray-drying, this approach could be potentially translated to a clinical setting for personalized medicine. In this scenario, the patient undergoes a surgery for removal of the tumor and the tumor cells can serve as a source of antigens for an individualized particulate vaccine, which can be administered therapeutically to avoid relapse [13]. Similarly, Crcarevska et al. [74] utilized the spray-drying technique (Buchi 190, Mini Spray Dryer) to encapsulate budesonide (anti-inflammatory glucocorticoid, poorly delivered at the injury site) into a cross-linked polyelectrolyte chitosan-alginate polymeric microparticle (coated with enterosolvent Eudragit) to suppress colonic mucosal injury and inflammation in a rat colitis model. As small microparticles (5.5. Hydrophobically modified hydroxyethyl cellulose has been used to functionalize PLGA nanoparticles to provide them with neutral charge and to evaluate their fate in mice after oral administration [89]. The study revealed that the smaller the nanoparticles, the greater the uptake by the Peyer’s patches. Also, the authors showed by confocal microscopy that the uptake of rhodamine-6G labeled nanoparticles by the Peyer’s patches was superior when using hydroxyethyl cellulose functionalized nanoparticles compared to positively charged-modified nanoparticles. Hydroxypropyl cellulose and hydroxyethyl cellulose have also been conjugated to fluoroquinolone-derived antibiotics for the delayed oral administration of these broad-spectrum antibiotics. In this regard, Amin et al. [90] showed that the oral administration of the nanoconjugates, and comparing to the effect of the free drug (i.e., ofloxacin) and to physical mixtures of the cellulose derivatives and the drug, resulted in a superior bioavailability for the preclinical model used. In the pharmacokinetic evaluation of concentration versus time, the areas under the curve from time zero to time infinity of the nanoconjugates were 2.1 and 2.3 times greater than those obtained for the free drug and for the corresponding physical mixture of ofloxacin with the cellulose derivative, respectively. Interestingly, a comparative study of the biodistribution profiles of a highly lipophilic drug (i.e., RR01) using poly(methacrylic acid-co-ethyl acrylate) nanoparticles and a standard formulation of the drug and hydroxypropyl cellulose (reference formulation)

Batch and microfluidic reactors in the synthesis of enteric drug carriers

after oral administration in dogs revealed a superior interindividual variability for the reference formulation [91]. Maximized bioavailability of the highly hydrophobic drug was reached when using the nanoformulation. 4.1.3 Polymers based on polyvinyl derivatives Polyvinyl derivatives, such as polyvinyl acetate phthalate, are commonly used as enteric coatings to provide with gastroprotection. As an example, an antifungal agent of low aqueous solubility and variable absorption and plasma concentration, itraconazole, has been micronized and gastroprotected with cellulose acetate phthalate and polyvinyl acetate phthalate to study the drug plasma concentrations in rats after oral administration [92]. Compared to a marketed form of the drug, the gastroprotected particles provided greater oral bioavailability, and higher concentrations of the drug were detected in the later sections of the tract due to the inhibition of drug precipitation. This crystallization inhibitory effect is also reported for polyvinyl acetate phthalate as coating on amorphous solid dispersions of celecoxib (a nonsteroidal anti-inflammatory drug) [93]. Polyvinyl alcohol (PVA) is commonly used to provide colloidal stability to nanoparticles due to its biocompatibility, biodegradability, neutral charge and hydrophilic character, having a large amount of hydroxyl groups prone to hydrogen bonding. Several nanoformulations intended for oral drug delivery make use of this synthetic polymer. For example, Ahlin at el [94]. included PVA in PLGA and polymethylmethacrylate (PMMA) nanoparticles for the oral delivery of enalaprilat (i.e., an angiotensin-converting enzyme inhibitor). PVA was used to control the size of the nanoparticles. The higher the concentration of PVA used, the smaller the nanoparticle size, due to the coalescence hindrance. PVA has also been reported as a crystallization inhibitor for several drugs (e.g., itraconazole [95], felodipine [96], etc.). Additionally, PVA has been used in combination with Eudragit polymers to modify the nanoparticle size and to tune drug release kinetics. In this regard, Bhadra et al. [22] formulated Eudragit L100-55 nanoparticles loaded with erythromycin stearate with different rations of PVA and evaluated drug release kinetics in vitro. The authors concluded that increasing PVA concentration rendered nanoparticles with reduced sizes, higher drug entrapment and a faster drug release. According to the authors, partially hydrolyzed PVA shows residual vinyl acetate groups, which have an amphiphilic character that favors polymer mixing and drug interaction. Polyvinyl acetate is also commonly used to stabilize nanoparticle suspensions. In fact, enhanced bioavailability has also been reported for furosemide nanosuspensions formulated with polyvinyl acetate when compared to the administration of the free drug. The pharmacodynamics effect of the drug, by evaluating the reduction in the systolic blood pressure, was also superior for the nanoformulation.

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4.2 pH-independent drug delivery Most orally administered carriers transit through the gastrointestinal tract with a limited control on the residence time, especially when intestinal motility is increased. It implies that, even if the carrier size was properly designed to control the drug release and promote the maximum bioavailability, the uncontrolled residence time could seriously affect drug concentration at specific sites of the intestine. Therefore, mucoadhesion has been used for a very long time in carriers designed to improve their residence time. Mucoadhesion can be achieved mainly by four types of interaction: (i) electrostatic (positive charged carriers), (ii) hydrophobic interactions, (iii) van der Waals interactions, and (iv) polymer chain interpenetration. Electrostatic attachment can be considered to attain a close interaction between the carrier and the epithelium, as well as to increase carrier and drug uptake via Peyer’s patches [97]. In this sense, electrostatic interactions between a positively-charged carrier surface, such as chitosan, and the negatively charged mucus (mucins) are widely explored. Besides chitosan, positively charged Eudragit polymers have been extensively used in the design of mucoadhesive polymeric carriers. As already mentioned, Eudragit polymers are synthetic polymers obtained by polymerization of acrylic acid, methacrylic acids and their esters. The physicochemical characteristics of Eudragit polymers, as well as their drug release, depend on the types of monomers and their proportion involved in the polymerization process. In the previous section, the pH-sensitive Eudragit types (L and S) were reviewed. However, there are also types of neutral Eudragit that are water insoluble, mucoadhesive, pH-insensitive in the gastrointestinal tract, swellable over the physiological pH, but still suitable for the sustained drug release in gastrointestinal tract. These Eudragit types are classified as [98]: Eudragit RL (highly permeable), Eudragit RS (low permeable), Eudragit NE (permeable), and Eudragit NM (permeable). Eudragit RL and RS are endowed with quaternary ammonium groups in the chloride salt form that impart a positive charge to the polymer. The dissociation of these groups in aqueous media controls the swellability and permeability of the polymers. The content of quaternary ammonium groups is higher in Eudragit RL than in RS, resulting Eudragit RL a more permeable polymer. Drug delivery in carriers made of Eudragit RL and Eudragit RS are directed by pore-diffusion. Alternatively, the drug release can be adjusted by using a mixture with different ratios of Eudragit RL/RS. Although these polymers are considered to be chemically stable, it was reported that several drugs (ibuprofen, diflunisal, flurbiprofen, and piroxicam) can interact with the ammonium groups presented in the polymer [99]. This interaction might limit drug loading efficiency, affecting its release profile and reproducibility [98, 99]. Eudragit NE and NM have similar properties to Eudragit RL/RS, but they are commercially available in aqueous dispersion with the emulsifier nonoxynol, and polyethylene glycol stearyl ether, respectively [98].

Batch and microfluidic reactors in the synthesis of enteric drug carriers

Clodronate loaded nanoparticles based on a cationic polymethacrylate Eudragit RL were tested in murine experimental colitis models, mimicking ulcerative colitis and Crohn’s disease [100]. Clodronate-loaded nanoparticles showed to be beneficial for the mitigation of experimental colitis, whereas the free drug did not exhibit any remarkable effect [100] (Fig. 6). Florfenicol, a synthetic broad-spectrum antibiotic was loaded in a blend of alginate and Eudragit RS [101]. Florfenicol release rate was decreasing as the pH was rising, which is desirable in a gastrointestinal tract drug release. When comparing with pure alginate microparticles, it was observed that approximately 40% and 20% of Florfenicol was released from alginate and alginate-Eudragit RS carriers at pH 7.4 in 1 h, respectively [101].

Fig. 6 Weight loss in the oxazolone (OXA)-colitis models (A) and in the 2,4,6-trinitrobenzenesulfonic acid (TNBS)-colitis model (B) during the whole treatment period (n ¼ 6, P < 0.05 compared with colitis control mice given saline, significance markers have not been added due to clarity reasons). (A): a distinct statistically significant mitigation of weight loss starting at day 2 could be detected considering clodronate loaded nanoparticles. (B): due to the severity of this colitis model and the low potency of the drug an effect was visible but non-significant. Reprinted with permission from Niebel W, Walkenbach K, Beduneau A, Pellequer Y, Lamprecht A. Nanoparticle-based clodronate delivery mitigates murine experimental colitis. J Control Release 2012;160(3):659–65. Copyright (2012) Elsevier.

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Proton pump inhibitors are commonly prescribed in the treatment of peptic ulcers and gastroesophageal reflux disease. However, some of them need to be gastroprotected when orally administered (e.g., lansoprazole). In this regard, the oil-in-water (o/w) emulsion solvent evaporation method was used for the synthesis of lansoprazole-loaded Eudragit RS100 nanoparticles [102]. Biodistribution analysis on ulcerated and nonulcerated regions of Wistar rats resulted in a preferential deposition of lansoprazoleloaded Eudragit RS100 nanoparticles in non-ulcerated areas. This can be rationalized by the high concentrations of positively charged proteins located at the ulcerated tissue [102]. Oral administration of lansoprazole loaded nanoparticles for 7 days healed 92.6–95.7% of gastric ulcers [102]. Heparin is an anticoagulant for the prevention of venous thrombosis and pulmonary embolism, with no oral bioavailability. Consequently, due to its short half-life and lack of oral absorption, heparin needs to be administered by the parenteral route under monitoring [103]. Oral formulations have resulted in limited success due to the short anticoagulant function exhibited. Polymeric nanoparticles based on a combination of PLGA and Eudragit RL yielded a high in vitro encapsulation efficiency of heparin (97%), due to the interaction of heparin with quaternary ammonium groups and the high content of these functional groups in Eudragit RL [103]. However, the best bioavailability in rabbits was obtained with Eudragit RL/Polycaprolactone NPs (20%), showing an anticoagulant activity for up to 7 h [103]. The most interesting fact of polymeric carriers is that they can be designed with a wide versatility just by selecting the proper polymeric combination to adjust their desired functions. Regarding the hydrophobicity, it was reported that a carrier composed of hydrophilic cellulose polymers was absorbed 100-fold less than the one composed with the hydrophobic cellulose polymer [104]. On the other hand, carriers synthesized from polymers such as poly(lactic acid) (PLA), PLGA poly(sebacic acid) (PSA), and poly(acrylic acid) (PAA) achieve mucoadhesion functionality by hydrogen bonding, as well as by hydrophobic interactions with the mucus. It was demonstrated that microcarriers composed of fumaric acid and sebacic acid exhibited higher retention time in rat gut compared to more weakly adhesive polymers such as alginate [105]. However, it should be considered that carriers could adhere nonspecifically to unintended locations or, even if they adhered, they could suffer a rapid clearance if they were in contact with loosely adherent mucus [97]. In addition to the variability in the gastrointestinal tract physiology associated with healthy individuals, gastrointestinal diseases also add complexity. For instance, inflamed tissues increase mucus secretion, affecting the viscosity of the mucus layer [97]. Targeted mucoadhesive carriers are designed with targeting ligands to enhance the binding specificity and decrease the elimination rate due to mucus turnover [97]. Lectins, invasins and Vitamin B12 represent the most extended mucus ligands. For instance, tomato lectin was conjugated in polystyrene particles, achieved an in vivo uptake increase of 50-fold compared with reference non-functionalized carriers [106]. In this case, the

Batch and microfluidic reactors in the synthesis of enteric drug carriers

lectin-conjugated carrier was preferentially uptaken via the villous tissue (15 times more than the uptake by the gut-associated lymphoid tissue) [106]. On the other hand, it was attempted the use of monoclonal antibodies and peptides as targeting moieties, as they have been shown to have high specificity in targeting and potential mucopenetrative properties [107]. Even though this approach can be successful by the parenteral route, it is challenging to achieve it by oral administration due the possible degradation of the monoclonal antibodies and peptides by exposure to acid pH and proteolytic enzymes. Polystyrene nanoparticles coated with anti-ICAM-1 antibodies expressed on the gastrointestinal epithelium and other cell types were produced [108]. Fluorescence and radioisotope tracing showed proximal accumulation, with preferential retention in the stomach, jejunum, and ileum. However, it was estimated that approximately 60% of the antibodies were enzymatically degraded [108].

5. Microfluidics for the synthesis of micro and nanocarriers in oral delivery The design and development of drug delivery vectors are complex, multi-step processes that require extensive efforts for chemical characterization, mass production, toxicity testing, and preclinical animal and clinical human trials. A plethora of micro and nanocarriers have been synthesized for oral drug delivery, including microparticles, microcapsules, microgels and nanoparticles made of various materials, including lipids, synthetic polymers, natural polymers and composite materials [109]. Conventional nanoparticle fabrication methods, such as nanoprecipitation and emulsification-evaporation, are largely associated with unstandardized, multi-step processes that need to be improved to assure the fabrication specifications, avoiding product batch-to-batch variations [110]. The scarce control on mixing conditions required to produce nanoparticles, especially as the reactor volume is increased, leads to polydisperse distributions and batch-tobatch variations. Considering that a successful (reliable and controllable) drug delivery system, as well as its inherent drug release profile, depends on the vector physicochemical properties (including the size, shape and composition), the use of additional steps to homogenize and sort the synthesized particles is commonly required when using batch-type reactors. Although oral drug delivery is the most widely used and most readily accepted administration route, it is challenging to fabricate carriers that afford achieving therapeutic drug levels due to the harsh conditions of the gastrointestinal tract, as well as the poor solubility, stability and bioavailability of many bioactive molecules and drugs. The ideal oral drug delivery carrier should fulfill the following items: (1) Protect the drug from premature degradation and release; (2) Release the drug at the desired site; (3) High drug loading; (4) Be produced using a reliable method with an excellent control on particle size and drug loading per particle; and (5) Load multiple therapeutic drugs to treat complex diseases.

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Microfluidics provides a wide variety of capabilities to manipulate fluids and ease a versatile production of nanoparticles with tunable size, shape and chemical composition. Micro-sized channels permit to handle fluid flows at the microscale very precisely, leading to highly reproducible synthesis and narrow size distributions in the resulting nanomaterials due to an efficient mixing driven by molecular diffusion. Microfluidic systems offer also some advantages against classical systems: homogenous reaction conditions, portability/easy wear, high sensitivity, low energy consumption, highly integrated multifunction and easy scalability [111]. In addition, the production of a single microfluidic system can be successfully parallelized to provide a reliable and reproducible mass production. Microdevices are manufactured by a serial or microfabrication steps such as thin-film deposition, photolithography, and etching. This procedure is followed for either the fabrication of the microdevice itself or the master mold (soft lithography). Conventional microfabrication techniques such as photolithography, wet/dry etching, and vapor deposition are based mainly on silicon/glass substrates, processed in cleanroom facilities. These requirements imply that the fabrication is expensive, laborintensive and not affordable for some users with a limited access to clean-room facilities. On the other hand, soft microfabrication requires inexpensive facilities and microdevices can be assembled at low-cost by users lacking proficiency in microfabrication. A complete description of the wide variety of lithographic techniques that can be used for the preparation of microfabricated devices and fundamentals of these methods can be found in previous reviews [112, 113]. Fig. 7 summarizes the most well-accepted microfluidic methods for the production of nanoparticles in drug delivery applications, mainly flow focusing, microvortices, template assembly, chaotic flow and droplet-based approaches [110]. Flow focusing relies on the hydrodynamic focusing of fluids, where the core fluid containing the samples of interest (drug, polymers, etc.) is sheathed by side fluids (Fig. 7). Two-dimensional flow focusing requires an easy fabrication of microfluidic systems, but those systems cannot successfully produce nanoparticles over a long period of time without channel fouling. Even though 3D flow focusing microdevices are complex to fabricate, their vertical and horizontal focusing flow patterns enable the isolation of the polymer-drug from the microchannel walls, avoiding channel blockage. Nanoprecipitation is an interesting approach for the synthesis of nanoparticles with small size ( 1 implies that there is a poor mixing in the reaction chamber and consequently a slow nucleation rate that will not result in ultrafine particles. Then, the reproducibility of the nanosuspensions is dependent on the mixing time of the solvent and anti-solvent solutions as compared to the time of nucleation and growth. On the other hand, the complex nature of precipitation results in an unfeasible scale-up from laboratory processes to industrial-scale batch processes. Considering these facts, it is well accepted that batch type reactors cannot provide the mixing control required, as it is only achieved by microfluidic reactors. Sylbin nanosuspensions, a therapeutic agent for liver diseases, were produced for oral application using a flow focusing approach with mixing by diffusion in a T-junction microchannel, and using acetone as antisolvent [115]. The sylbin particle size depends on the supersaturation achieved by the antisolvent/solvent ratio (the volume of antisolvent controls the nucleation and crystallization kinetics), the drug concentration, as well as the residence time. Surprisingly, the smallest nanoparticles (39 nm) were not achieved at the smallest residence time [115]. If the flow rate of the solvent and antisolvent were kept constant, the diffusion and transfer between the drug solution and antisolvent could not be completed, and the drug particle size evolved in the collecting vial once the mixing is achieved by diffusion. Core/Shell nanoparticles for oral delivery of chemotherapeutics to treat colorectal cancer were produced using a two stage microfluidic system [116]. The core was produced in a cross-junction flow focusing microfluidic device by controlling the mixing of a hydrophobically modified chitosan derivative (modified with N-palmitoyl chitosan) with a Paclitaxel stream, focused with sheath flows of water at basic pH. N-palmitoyl chitosan chains enabled to improve the loading of paclitaxel and the nanoparticle selfassembly. The resulting monodisperse nanoparticles were downstream coated in a Tesla micromixer, using a pH-sensitive polymer (Eudragit) to protect the nanoparticles from acidic pH [116]. The Tesla micromixer uses the Coanda effect, where the fluid wets the nearby surface and remains wetting it even when the streamlines modify their direction, to efficiently mix the inlet streams by a kind of chaotic flow (Fig. 7). At the inlet of the Tesla micromixers, part of the fluid is deflected due to Coanda effect, while the remaining fluid flows through the curved-section until they finally merge [116]. These flow disruptive patterns enable a fast mixing of the reagents. In a recent study, a continuous crystallization process was successfully performed at high pressure to decrease the mixing time of reagents in an impinging jet chamber

Batch and microfluidic reactors in the synthesis of enteric drug carriers

(residence time, 300 ms) [117]. Four different poorly water-soluble drugs (Class 2 according to Fig. 2), phenytoin, miconazole, bezafibrate, and flurbiprofen, were nanoprecipitated using ethanol and water as solvent and anti-solvent, respectively [117]. After purification, it was observed that the particle size and size distributions were smaller and more monodisperse than those prepared with a conventional nanoprecipitation process in a glass beaker, resulting in superior in vivo absorption [117]. Microvortices-based microfluidic chips were generated to overcome the flow inefficiencies of slow diffusive mixing, promoting the achievement of 3D chemical profiles, and assuring a fast and efficient micromixing (Fig. 7). This ability of rapidly mixing reagents provides a homogeneous reaction environment and higher productivity of nanoparticles than diffusive mixing-based nanoparticle synthesis [118]. Microvortices in a co-flow microfluidic device were used to promote the micromixing of a lipidcontaining organic phase with an aqueous buffer, resulting in the self-assembly of lipid bilayers [119]. This approach was considered to boost bacteriophage therapy to treat gastrointestinal infections. Orally delivered phages are not very efficient in treating infections, as these tend to have short residence times in the gastrointestinal tract due to clinical symptoms such as diarrhea [119]. In addition, phages can be inactivated in the stomach acidic conditions, resulting in low doses of phages delivered downstream in the gastrointestinal tract. Escherichia coli T3 podovirus and myovirus Staphylococcus aureus phage K were loaded in a continuous flow into liposomes with a particle size ranging between 100 and 300 nm [119]. The lipid dispersed in an isopropanol alcohol stream was sheathed by a coaxial aqueous stream, resulting in an excellent mixing by microvortices across fluid interfaces, and in the continuous assembly of liposomes of narrow and controlled size distributions. Although the liposome production was controlled, phages could still bind to external lipid bilayers, decreasing the yield of encapsulated phages [119] (externally bound phages are easily inactivated in stomach acidic conditions). Droplet-based microfluidics has been demonstrated to be one of the most efficient microfluidic approaches due to the easy control of discrete volumes of immiscible fluids (Fig. 7). Two types of flow focusing devices have been considered in the droplet-based microfluidics approach to produce spherical vectors: (1) a chip-based flow-focusing device (T or Ψ shaped microchannel integrated on a chip), and (2) a capillary-based flow-focusing device (coaxially aligned microcapillaries). The droplet size is tuned by adjusting the flow rates and ratios of the fluids, the viscosity, and the microchannel geometry. Finally, the droplet can be solidified inline downstream by chemical or physical methods. For instance, hardened liquid droplets could be achieved by polycondensation [120], radical polymerization [121], ionic crosslinking [122], thermosetting [123] and solvent evaporation or extraction [124]. Although this microfluidic pattern is robust, it suffers from the inability to produce nanoscale drug carriers and low production yields. The type of droplet that can be produced using this approach can be classified as: (i) Single emulsion, (ii) Double emulsion, and (iii) Janus. Single emulsion implies the

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emulsification of one fluid in a second immiscible fluid, including oil-in water (o/w) or water-in-oil (w/o) types. The double emulsion type is derived from the combination of two single emulsion units. This type of droplet requires the use of, at least, three different fluids, i.e., water-in-oil-in-water (w/o/w) or oil-in-water-in-oil (o/w/o). Finally, Janus particles combine two or more distinct materials with different physical properties together. An off-the-shelf capillary based co-axial drop-flow microfluidic device was selected to produce acrylate based polymerizable ketoprofen-loaded droplets [125]. The droplets were cured downstream by UV light, without damaging the encapsulated ketoprofen. As a result, microbeads in the size range of 200–380 μm and with high encapsulation efficiency (80–100%) were achieved [125]. At pH 1.2, the amount of ketoprofen released within the residence time in the stomach was less than 4%, whereas at pH 6.8 a gradual ketoprofen release was observed, tuning the % release as a function of the monomer (ethyl acrylate) concentration used [125]. Dual drug delivery harnesses the synergistic drug effect, providing significant advances in the treatment of many complex diseases [126]. The main hurdle in dual drug therapy is the difficulty in co-encapsulating hydrophobic and hydrophilic drugs together in the same carrier, since drugs with different polarities have different solubilities. Core-shell polymeric particles containing a hydrophobic model drug in the core and a hydrophilic model drug in the shell have been produced via UV initiated free radical polymerization of double droplets. On a study by Khan et al. [127], droplets were produced in continuous flow by two co-axial capillary-based drop-flow microfluidic devices. It is known that ketoprofen uncontrolled release from conventional dosage forms can cause serious systemic and gastric problems [128]. However, the release of ranitidine HCl at acidic pH (stomach) could suppress the gastric irritation effects associated with nonsteroidal anti-inflammatory drugs like ketoprofen [127]. Therefore, ketoprofen was loaded in the poly(methyl acrylate) core, whereas ranitidine HCl was loaded at the poly(acrylamide) shell. Core-shell microparticles revealed minimal ketoprofen release at low pH (1.2), and a maximum release at the colon pH (7). Interestingly, the release of each drug was tuned by changing the dimensions of the core and the shell [127]. The same microfluidic approach was used to produce poly(acrylamide) Trojan microparticles containing ketoprofen-loaded poly(ethyl acrylate) or poly(methyl acrylate) nanoparticles [129]. At pH 6.8, Trojan microparticles released 35% of the encapsulated ketoprofen-loaded nanoparticles over a 24-h period. Glucagon-like peptide-1 (GLP-1) and a dipeptidyl peptidase-4 inhibitor (iDPP4) drug were successfully combined in a single carrier to face the side effects associated with each drug separately [130, 131]. GLP-1 is one of the most promising drugs to treat type 2 diabetes mellitus, but it is degraded by the DPP4 enzyme in less than 2 minutes [132]. A single-emulsion drop flow approach was used to produce monodisperse microparticles with high encapsulation efficiency [130, 131] (Fig. 8). GLP-1 was first loaded into

Fig. 8 (A) Schematic representation of the drop-flow microfluidic approach used to produce the pH-responsive micropaticles, coloaded with GLP-1 and DPP4 inhibitor. (B) SEM images of microparticles at different pHs, the enteric coating was degraded at a pH value above 6. Reprinted (adapted) with permission from Araújo F, Shrestha N, Shahbazi M-A, Liu D, Herranz-Blanco B, Ma€kila€ EM, et al. Microfluidic assembly of a multifunctional tailorable composite system designed for site specific combined oral delivery of peptide drugs. ACS Nano 2015;9(8):8291–302. Copyright (2015) American Chemical Society.

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nanoparticles composed of PLGA or porous silicon. These nanoparticles were further sequentially modified with chitosan, a mucoadhesive polymer, and with an oligoarginine (cell-penetrating peptide), to increase the permeability of nanoparticles across the intestinal cells [131]. The nanoparticles and the iDPP4 were encapsulated within an enteric polymer HPMC acetyl succinate [131]. This formulation inhibited the premature release of GLP-1 and its degradation in vitro in gastrointestinal tract conditions, conferring pH-sensitivity to the microparticles, and thus enabling the release of drugs in simulated intestinal conditions (Fig. 8). The GLP-1 loaded, pH-sensitive PLGA carrier was further evaluated in vivo in a non-obese type 2 diabetes mellitus rat model, achieving a hypoglycemic decrease of 44% [130]. In addition, the dual release of GLP-1 and iDPP4 enabled to achieve an enhancement of the plasmatic insulin levels 6 h after oral administration [130]. A hierarchically structured carrier was also designed using the single-emulsion drop flow approach [133]. In this case, porous silicon nanoparticles were functionalized with hyaluronic acid. Acorbyl palmitate (hydrogel promotor) and budesonide (a glucocorticoid for IBD therapy) were co-encapsulated in the mesopores of functionalized porous silicon nanoparticles. Finally, a hierarchically structured microparticle was produced in continuous drop flow using HPMC acetate succinate [133]. An in vitro budesonide delivery test was compared between the hierarchical microparticles and the free drug, resulting in a major reduction in systemic budesonide release, and consequently the reduction of possible side effects. In vivo studies also confirmed the efficient protection of the enteric polymeric shell, as well as the local drug delivery of microparticles to inflamed sites of the intestine in an IBD model. Nanoparticles selectively targeted the inflamed sites of intestine (positive charged) by electrostatic interaction, and locally released the drug in response to inflammation, and in a gradual manner, over a prolonged duration [133]. A similar concept was explored to load atorvastatin and celecoxib in halloysite nanotubes. These two model drugs act synergistically on colon cancer and its prevention [134]. The pH-responsive HPMC acetate succinate polymer was selected to produce pH sensitive microparticles by the aforementioned single-emulsion drop flow approach (Fig. 7) [134]. In a different study, the same flow-focusing approach and pH-responsive polymer were used to produce pH-responsive microparticles by a double emulsion of halloysite nanotubes, further modified with the mucoadhesive polymer poly(methyl vinyl ether-comaleic) and loaded with curcumin [135]. A droplet-based microfluidic flow device was used to produce hollow protein (zein) microcapsules for applications in gastro-intestinal drug delivery [136]. Particle size, internal structures and permeability were easily tuned by varying the disperse phase flow rate and zein concentration [136]. Liposome vesicles with high uniformity in both size and encapsulation efficiency (94%) have also been produced by glass capillary microfluidics using a drop flow approach with a w/o/w double emulsion [137]. In this case, the liposome vesicle w/o/w template consisted of a shell, composed of liquid unsaturated phospholipids and saturated powdered phospholipids. This composition was selected to

Batch and microfluidic reactors in the synthesis of enteric drug carriers

overcome the low transition temperature of unsaturated powdered phospholipids [137]. The core was composed of the hydrophilic active doxorubicin hydrochloride, either alone or with nanomagnetite, for potential application as magnetic resonance imaging contrast agent for biomedical diagnosis. The final phospholipid vesicles were formed after a solvent removal step by dewetting. The doxorubicin-loaded phospholipid vesicles showed better sustained release than free DOX aqueous solution, and were considered potential in vivo carriers for improved oral bioavailability [137]. Nanoprecipitation has also been considered to produce pH-sensitive hypromellose acetate succinate microdroplets by the drop flow approach [138]. For this purpose, the droplets were loaded with porous silicon nanoparticles, previously conjugated with the Fc fragment of immunoglobulin G, loaded with Glucagon-like peptide-1, and coated with the mucoadhesive chitosan [138]. The supersaturation achieved by using the proper solvent (acetone) and antisolvent (1% PVA, pH 3.7), enabled precipitation of the hypromellose acetate succinate and, subsequently, the entrapment of the aforementioned GLP-1 loaded porous silicon nanoparticles. The pH-responsive capacity of the microparticles allowed for the protection of the payload at low pH values (gastric pH) and triggered a controlled drug delivery at intestinal pH, with enhanced GLP-1 absorption across the intestinal monolayers [138]. A gas-liquid segmented flow microfluidic approach, a variant of the drop flow approach, was selected to manufacture, in continuous flow, polymeric cellular dosage forms [139]. An extruder injects the melting polymer through a capillary channel, where gas slugs are injected at fixed time intervals to control the size of the dosage forms. These polymeric cellular dosage forms float over the gastric content, providing a longer and more predictable gastric residence time [139]. Flow lithography is the microfluidic technique that combines photolithography with microfluidics (Fig. 7). This technique allows for the controlled fabrication of microparticles, where a pattern is exposed in photopatternable polymers into a microchannel and under a pulsed laminar flow. This approach is limited to optically transparent polymers, since the UV light should penetrate through them. On the other hand, it can only be used for non-photosensitive drug molecules, to avoid the UV-light induced damage. Against previous techniques, flow lithography is the only approach available to produce non-spherical shaped particles, which could have a difference cellular internalization and intracellular trafficking, when compared to their spherical-shaped counterparts [140]. In addition, it is rationalized that oral microdelivery particles may be designed to be flat (disc shaped), due to the higher contact area with the intestinal lining [141]. Chaotic advection is a microfluidic technique for enhancing mixing efficiency. It utilizes flow disruptive patterns that passively mix fluids within the cross section of a microchannel (Fig. 7). The staggered herringbone and Tesla patterns are the most widely used microfabricated geometries to achieve chaotic advection, allowing for the creation of a fast diffusional mixing when compared with hydrodynamic flow focusing approaches.

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Bulk nanoprecipitation of polymer mixtures (amphiphilic, hydrophobic and hydrophilic) is challenging due to the unequal distribution of the polymers along the reactor (polymer solubility), as well as mixing limitations. These issues have been addressed using a chaotic mixer featuring staggered herringbone structures. This type of approach resulted in the production of multicomponent polymeric nanoparticles with a diminishment of both particle size and polydispersity index [142].

6. Conclusions To obtain reproducible drug release profiles from particulate carriers, the production of monodispersed drug loaded particles is required, which is very challenging when using conventional discontinuous production methods due to their mixing and mass transfer limitations. The mass scale production to reach clinical translation of nanotechnologybased therapies is still challenging. Microfluidic technologies endow biomedicine with a plethora of tools to control the production of active pharmaceutical ingredients loaded in nano- and micro-scaled carriers. This technology is gaining a tremendous relevance and its ongoing sophistication holds the promise of facing the current challenges of orally administered drugs, by which carriers can be engineered to increase the efficacy of gastrointestinal tract delivery. The properties of the synthesized drug carriers using microfluidic platforms can be effectively and reproducibly modified by tuning flow rates and geometrical features. The nano- and micro-encapsulation of active principles provides with demonstrated benefits over the free administration of the drug in many settings. However, there are still tremendous limitations in their clinical translation and only a few nanomedicines have reached the market. In targeted therapies, the use of homing molecules still lacks a demonstrated efficacy in vivo, also immune recognition and their large-scale production are still barriers to overcome, which represents a limitation for this technology. Microfluidics can assist in their large-scale production and, more importantly, in obtaining materials that cannot be produced by using conventional batch-type reactors.

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3D printing in oral drug delivery Natalja Genina, Johan Peter Boetker, Jukka Rantanen

Department of Pharmacy, Faculty of Health and Medical Sciences, University of Copenhagen, Copenhagen, Denmark

1. Oral delivery 1.1 Strategies to overcome physiological limitations of the gastrointestinal (GI) tract It is well-known that the drug delivery systems (DDS) administered orally will be exposed to the different physiological conditions before a complete or partial absorption occurs. This would include (i) the change in the pH of the gastrointestinal tract (GI) between the oral cavity, stomach, small intestine, and colon; (ii) varying gastric and intestinal transit time; (iii) variable enzymatic activity at the membranes of the GI tract; (iv) food intake, and (v) severity of the disease. Before absorption can occur, the active pharmaceutical ingredient (API) has to be dissolved and/or suspended in the GI liquid. One of the key challenges is that the solubility of the most active pharmaceutical ingredients (APIs) depends on the pH. Therefore, the most common strategy is to formulate the DDS in such a way that the release of the API occurs in the part of the GI tract that has the optimal pH. For example, a drug that has the highest solubility in an acidic environment may be aimed to be released and at least partially absorbed in the stomach. In some cases, it would be preferred to achieve a long-term release of the active compound in the stomach. For instance, it could be crucial to minimize the frequency of a dosage form intake in order to improve treatment adherence. Furthermore, a long-term residence time of the dosage form in the stomach may be needed to treat diseases located in the gastric region. There is, however, a physiological limitation in regards to the gastric transit time being in the magnitude of 5 min to 2 h. This obstacle can be overcome by designing gastric retentive systems, such as floating dosage forms, unfolding polymer sheets, swelling hydrogel balloons, etc., that would not be able to pass through the pylorus and thus stay in the stomach until they are dissolved, disintegrated, and/or eroded to the size that would fit the aperture. An interesting approach was suggested by Bellinger et al. [1], where an oral capsule shell first dissolves in the stomach and subsequently deploys a star-shaped dosage form. The developed DDS could retain and release the anti-malaria drug in the stomach for up to 14 days in pig models (Fig. 1). The star shaped DDS contains regions that are able to dissolve in near neutral pH and allow the remaining elements safe pass through the GI tract. In this work, additive manufacturing, e.g., 3D printing, was used to print positive molds in order to fabricate polydimethylsiloxane Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00009-0

© 2020 Elsevier Inc. All rights reserved.

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Fig. 1 (A) A photograph of the dosage form encapsulated into a 00el hard gelatin capsule, (B) a photograph of the deployed dosage form, (C) a schematic representation of deployment of the encapsulated dosage form in the gastrointestinal tract. (Reprinted with permission from the American Association for the Advancement of Science (AAAS): Bellinger AM, Jafari M, Grant TM, Zhang S, Slater HC, Wenger EA, Mo S, Lee Y-AL, Mazdiyasni H, Kogan L, Barman R, Cleveland C, Booth L, Bensel T, Minahan D, Hurowitz HM, Tai T, Daily J, Nikolic B, Wood L, Eckhoff PA, Langer R, Traverso G. Oral, ultralong-lasting drug delivery: application toward malaria elimination goals. Sci Transl Med 2016;8:365ra157.)

(PDMS) negative molds for melt-molding of the drug-containing stellate geometries. Research has also been conducted to design floating DDS that resist gastric emptying by floating on the stomach contents. This has been previously achieved by manufacturing (i) hydrodynamically balanced systems [2]; (ii) carbon dioxide-generating systems, and (iii) freeze-dried systems. It is evident that the inclusion of the gas-filled voids within the dosage form makes it lighter than the contents of the stomach, resisting the housekeeping function of the pylorus. However, the possibility of flexible adjustments of the internal structure of the dosage forms by conventional manufacturing methods is limited. Therefore, 3D printing was used to fabricate the floating drug-containing dosage forms with the unique porous internal structure [3]. Currently, injection is the most used method for peptide drug delivery as the traditional oral route of administration represents inherent limitations of poor absorption and degradation of these macromolecules in the GI tract. However, there is a constant demand to design biopharmaceuticals for oral drug delivery due to patients’ inconvenience imposed by daily injections. Recently, the Danish company Novo Nordisk has received the Food and Drug Administration (FDA) approval for the first glucagon-like peptide-1 (GLP-1) analog semaglutide in a tablet form (Rybelsus®). A key element in the functionality of this first oral product is the absorption enhancer, and additive manufacturing opens new design principles when contrasting with compaction-based technologies. An important benefit of using additive manufacturing is the ability to design labile API-containing dosage forms that would avoid harsh gastric and intestinal conditions and the first pass metabolism.

3D printing in oral drug delivery

The oral cavity is also a potential drug delivery route, because of the mild pH conditions and the reduced enzymatic activity. For example, Minirin®, a peptide (desmopressin)containing freeze-dried drug product, intended for sublingual administration, exists on the market. This route of administration would, among other benefits, give a fast onset of action that can be critical for some disease conditions. Besides biomolecules, other labile drugs could have an increased systemic bioavailability and a rapid onset of action via administration through the oral cavity. For example, naturally derived antichlamydial agent Biochanin A undergoes an extensive first pass metabolism, and buccal/sublingual dosage forms, e.g., films, could be an option to avoid that [4]. Furthermore, the oral delivery of the API can be useful to treat local (oral) inflammation. In this regard, wearable oral drug delivery devices such as a mouthguard are an attractive alternative to the conventional oral dosage forms [5]. 3D printing is one of the advancing techniques that can produce on-demand patient-tailored mouthguards with customizable design and the desired release rate of the API (Fig. 2). Potentially, the systemic delivery of various active compounds is possible with this wearable oral drug delivery device. Though, the removal and cleaning procedure should be considered if used multiple times. Increased bioavailability of peptide formulations can be achieved by peroral administration of a drug delivery system that is self-oriented in the stomach and injects the peptide (insulin) through the gastric mucosa without perforation of the tissue [6]. The device resembles a roly-poly toy, where the bottom part consists of a high density material, e.g., stainless steel, and the upper part is made of a low-density material such as polycaprolactone (PCL) (Fig. 3). The peptide-containing millipost is attached to the spring that is fixed with caramelized sugar in the top of the device. Upon hydration, the sugar dissolves and thus releases the applicator. It was shown that this device delivers peptide plasma levels close to the ones obtained from subcutaneous injections. The authors do not specify the methods used to produce the components of the shell for this device, but 3D printing can be well-used to fabricate those parts. Despite the advances in the delivery of the drugs through the oral and gastric mucosa, the majority of the APIs released from DDS are still aimed to be absorbed in the small intestine. The reasons for this are (i) increased surface area as compared to the stomach, (ii) relatively long and relatively constant transit time of around 3 h, and (iii) mild environment with pH from 5 to 7. To avoid the direct contact with the acidic environment of the stomach that would be detrimental for acid-labile APIs, a gastroresistant coating is commonly applied to the dosage forms. The polymers used for the enterocoating start dissolving only when the pH increases above the pH value of the stomach, i.e., when the dosage form leaves the stomach and enters the small intestine. 3D printing can be used as well as the other conventional methods to provide such protection for the acid-labile compounds. Furthermore, 3D printing has been used in a similar manner as described by Bellinger et al. [1] to fabricate a luminal unfolding microneedle injector (LUMI) for oral delivery of macromolecules to the small intestine [7].

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3D manufacture

Compound + polymers

Tunable release

Hot melt extrusion

Drug release

Data acquisition

Filament 1

Filament 2 Time

Intraoral scan 3D-printed mouthguards Horizontally sliced

Maxillary anatomy

Vertically sliced

Customizable design

FDM-based 3D printing

Evaluation in humans

Fig. 2 The steps involved in the manufacturing of wearable personalized drug delivery mouthguards by 3D printing. (Reprinted with permission from AAAS: Liang K, Carmone S, Brambilla D, Leroux J-C. 3D printing of a wearable personalized oral delivery device: a first-in-human study. Sci Adv 2018;4:eaat2544. https://doi.org/10.1126/sciadv.aat2544)

3D printing in oral drug delivery

Fig. 3 (A) A photograph of an ingestible self-orienting millimeter-scale applicator (SOMA) and (B) a schematic illustration of the internal structure of SOMA. (Reprinted with permission from AAAS: Abramson A, Caffarel-Salvador E, Khang M, Dellal D, Silverstein D, Gao Y, Frederiksen MR, Vegge A, Hubálek F, Water JJ, Friderichsen AV, Fels J, Kirk RK, Cleveland C, Collins J, Tamang S, Hayward A, Landh T, Buckley ST, Roxhed N, Rahbek U, Langer R, Traverso G. An ingestible self-orienting system for oral delivery of macromolecules. Science 2019;363:611–5. https://doi.org/10.1126/science.aau2277.)

Colonic drug delivery is of importance for the treatment of colonic diseases such as Crohn’s disease, colon cancer or ulcerative colitis, and holds potential for protein delivery. This region of the GI tract has a pH of 6–7.5, and the colonic transit time is between 2 and 48 h. Again, enterocoating of the dosage forms with the polymers that start dissolving in the colonic environment is the most common strategy to target this part of the GI tract. The coating, which is usually performed in coating pans or fluidized bed coaters, is an extra manufacturing step with an extra financial cost. Linares et al. [8] have proposed to combine 3D printing with injection volume filling in a single step to produce the dosage forms for colon-specific delivery. Oral drug administration includes the transition of the dosage form from the oral cavity to the stomach via the esophagus by the act of swallowing. This action can represent severe challenges for people, who possess swallowing difficulties (e.g., children and elderly people). The main reason for such difficulties is the inappropriate size, shape and/or surface characteristics of the solid dosage form. Manipulating the physical dimensions, selecting the easy-to-swallow shape (e.g., almond shape), and/or polymeric coating of the DDS to decrease surface roughness are the common strategies to minimize dysphagia. Administration of orodispersible formulations is another alternative to avoid difficulties in swallowing [9]. The first ever FDA approved 3D printed drug product, Spritam®, is an orodispersible dosage form, containing either 250, 500, 750, or 1000 mg of the antiepileptic drug levetiracetam [10, 11]. The porous tablet disintegrates in the mouth with a sip of liquid within seconds. It would be extremely difficult to compress such a high dose orodispersible tablet using the conventional tablet manufacturing methods, highlighting the unique possibilities provided by 3D printing. Besides orodispersible tablets, orodispersible films (ODF) can be the option to minimize the swallowing difficulties [9]. An ODF is

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usually a stamp-size drug-containing polymeric DDS that is fabricated by solvent casting onto large impermeable liners. The required size of the film is then cut out to yield the desired dose per dosage form. In this regard, 3D printing can be used for on-demand fabrication of an ODF with the desired physical dimensions. The layer-by-layer printing results in a more porous internal structure of the dosage that in contrast to solvent casting can result in a faster dissolution rate of the API from the 3D printed ODF [12]. Solvent casted ODF lack a good physical and chemical stability of the API on storage, and ODF are very much affected by changes in the ambient conditions, especially the relative humidity. Therefore, expensive packaging is provided for each ODF. In addition, any variation in the thickness of the film during production may have an effect on the dose precision of the final dosage forms. The fabrication of placebo ODF films in advance and printing of the precise dose of API onto them on-demand can be a way to minimize the problems related to the dose accuracy and storage stability of the solvent casted ODF [13]. Fig. 4 is summarizing the target areas in GI tract with selected examples of the innovative 3D printed DDS for each part of the GI tract.

Fig. 4 Illustration of the target areas in GI tract with selected examples of the innovative 3D printed DDS for each part of the GI tract. (The image of a complete digestive tract was taken from Servier Medical Art [WWW Document], 2019. https://smart.servier.com/smart_image/complete-digestive-apparatus-3/ [Accessed 29 October 2019].)

3D printing in oral drug delivery

1.2 Introduction to the different release mechanisms Drug release from the dosage form is one of the most, if not the most important critical quality attribute of the drug product. The drug can be released in several ways from the oral solid dosage form, depending on the composition of the formulation, presence of the release-limiting barriers, the internal structure of the dosage unit, chemical binding to ion-exchange resins, and incorporation in the osmotic pump. Though, the oral modified-release dosage forms can be mainly considered as (i) monolithic or matrix systems and (ii) reservoir or membrane-controlled systems. In the first case, the API is dissolved/dispersed in hydrophilic/lipid/insoluble polymer matrices, and the drug is released as the matrix dissolves/swells/erodes. In the case of an insoluble or non-erodible matrix, the solvent penetrates into the matrix and dissolves the drug, and the drug is released through the pores by diffusion. In the latter case, the membrane (coating) determines the release rate of the drug. Osmotic pump systems belong to this group, where a semi-permeable membrane controls the diffusion rate of water inside the systems and by that determines the release rate of the drug through the hole made in the membrane. For both types, the selected polymer (lipid) and its physicochemical properties as the matrix or membrane former would define the release rate of the drug from the DDS. Polymethacrylates (Eudragit®) and cellulose derivatives are widely used polymers in those systems. Tremendous work has been done to explore the suitability of these polymers and other pharmaceutically approved polymers together with plasticizers for the 3D printing processes to produce modified-release peroral dosage forms [14, 15] that would replicate the existing modified-release dosage units [16]. There is a variety of terms used to describe the different ways of the drug release. Immediate release indicates that the drug is released immediately following administration (e.g., uncoated tablets); delayed release indicates that the drug is not released immediately, but at a later time (e.g., enterocoated and pulsatile released dosage forms); controlled release indicates the release of the drug at a constant rate over a long period (e.g., osmotic pump); sustained release means an initial release of the therapeutic dose and then a gradual release of the drug over an extended period (e.g., matrix systems). Furthermore, the sustained release of the drug can be achieved by incorporation of multi-particulate systems (e.g., pellets) in a single dosage unit (e.g., capsule), where one portion of the pellets is uncoated to provide an immediate release, whereas another one is membrane-coated to provide a gradual release. The thickness and the type of coating can be different to guarantee the steady release of the drug over a long time. Ritalin® is the commercial product for the treatment of attention deficit hyperactivity disorder and narcolepsy that contains the multi-particulate systems with varying release properties. Despite the evident benefits of such dosage unit, coating the pellets with different membranes and filling of the capsule with different types of pellets is complex. In this regard, 3D printing can be considered as a viable approach for producing modified release dosage forms in a single production step by combining different polymers in a single dosage unit [15]. Furthermore, the release rate of the drug from the DDS can be adjusted by 3D printing of either

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the drug or material gradient within the dosage form, again in a single step that would, in turn, modify the inner structure of the dosage unit [17]. The drug release can also be modified by altering the solid content (infill parameter) of the dosage form, and by that modifying the available surface area for release of the API [18]. This is done by a simple manipulation of the process parameters in the digital file of the object to be printed, without the need of adjusting the formulation [19] (Fig. 5). In addition, the drug release from the dosage form can be modified by manufacturing different tablet shapes of different geometries. For instance, a toroidal shape of an oral dosage form would give zero-order kinetics [20]. In this context, 3D printing would be the method of choice, as it would be challenging to produce complex geometries by, for example, powder compaction [20]. The versatility of 3D printing can be used to fabricate the molds, both positive and negatives, for casting, compressing or molding of the dosage forms with unique design, and consequently release rate [21] (Fig. 6). Furthermore, a unique geometry of the molds can be used to produce tablets with a programmed disintegration profile, by varying the thickness of the different regions in the dosage form [22].

Fig. 5 Photographs of 3D printed honeycomb-like tablets with different cell sizes. (Reprinted with permission from Elsevier: Kyobula M, Adedeji A, Alexander MR, Saleh E, Wildman R, Ashcroft I, Gellert PR, Roberts CJ. 3D inkjet printing of tablets exploiting bespoke complex geometries for controlled and tuneable drug release. J Control Release 2017;261:207–15. https://doi.org/10.1016/j.jconrel.2017.06.025.)

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1.3 Fixed dose combination drug therapy (FDCDT) There are several reasons for formulating the medicine as a combination therapy product. First of all, the fixed dose combination can significantly reduce the side effects and improve the efficacy of the therapy by reducing the dose of the APIs. For instance, antiretroviral Kaletra® contains both lopinavir and ritonavir, where a low dose of ritonavir acts as a booster of lopinavir, enhancing its effect.

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Secondly, patients’ adherence to the therapy can be significantly improved by minimizing the number of different dosage units taken daily [23]. The treatment of cardiovascular diseases usually requires simultaneous administration of multiple drugs. Polycap™, a capsule that contains five different APIs (granules), was designed to reduce the number of different units required each day. The possibility of 3D printing a polypill similar to Polycap™ was successfully employed [24]. The 3D printed tablet consisted of five compartments where: two of the compartments provided an immediate release of aspirin and hydrocholorthiazide, and the other three compartments delivered a sustained release of pravastatin, atenolol, and ramipril (Fig. 7). The modified release of each drug was achieved by the utilized polymers in the formulation. The potential benefit of 3D printing is that a five-in-one tablet was done using a single 3D printer with separate ink cartridges for each API, meaning that simultaneous printing of different APIs is possible. This manufacturing approach holds promises for on-demand tailoring of a particular drug combination for the need of each patient. Finally, this combination strategy can be beneficial to prevent the development of drug resistance. For instance, the World Health Organization (WHO) recommends FDCDT such as the oral administration of rifampicin and isoniazid for an effective management of tuberculosis, and to avoid the transmission of the disease. Interestingly, simultaneous administration of both drugs as immediate release formulations would be undesired, because rifampicin is unstable in the presence of the isoniazid in the acidic environment, even though it is efficiently absorbed in the stomach [25]. The gastroretentive combination system was fabricated to allow a long-term release of rifampicin in the stomach, while the same dosage unit contained an enterocoated capsule with isoniazid [26]. In this context, the potential of 3D printing, such as the ability to produce the system with a unique internal structure, was employed to print a dual-compartmental

Fig. 7 (A) A schematic illustration of a polypill with five active pharmaceutical compounds incorporated in a single dosage form, and (B) photographs of the polypill. (Reprinted with permission from Elsevier: Khaled SA, Burley JC, Alexander MR, Yang J, Roberts CJ. 3D printing of five-inone dose combination polypill with defined immediate and sustained release profiles. J Control Release 2015;217:308–14. https://doi.org/10.1016/j.jconrel.2015.09.028.)

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dosage unit [27]. The compartmental system was designed to physically isolate both APIs from each other, and furthermore, the isoniazid-containing compartment was sealed to delay the release of the API. The similar principals of isolation and unidirectional release could be used to deliver encapsulated biologics and human microbiota to treat certain diseases more efficiently. 3D printing can also be used to design multi-compartmental systems that hold both liquid and solid phase APIs, or multiple liquid phase APIs in different compartments, in a single dosage unit [28]. Such design of the dosage forms would be extraordinarily challenging to make with the existing, conventional manufacturing solutions.

2. Personalized medicine 2.1 Motivation for the patient-tailored medicine The current disease management is based on a ‘one size fits all’ principle. This means that only standardized doses are available on market, which derived from the needs of the majority of the patients. However, the minority of the patients lacks the correct dose. The administration of suboptimal doses may result in severe side effects or underdosing with no therapeutic effect. To fulfill the need of each patient, the personalization of the dosage forms has been proposed [29].

2.2 Prerequisites for the production of personalized medicine Patient-tailored therapy can be defined based on the weight, gender, age and metabolizing capacity of the individual. For instance, the required therapeutic dose for a child (e.g., paracetamol dose) is mainly determined by the weight of the child. Nowadays it is becoming more common to use diagnostic tests to look for disease-specific and/or treatment-specific biomarkers (e.g., genes), to select the targeted therapy [30]. The continuous monitoring of the predictive biomarkers helps to optimize and on-demand adjust the treatment regimen, and forecast the probability of a positive outcome from a given therapeutic treatment [31]. Besides pharmacogenomics and pharmacogenetics, noninvasive continuous monitoring of other parameters can be beneficial to get the status of the patient’s health at any time point. Wearable sensor arrays for in situ analysis of different parameters in sweat can be a useful tool for on-demand adjustment of the ongoing therapy [32]. Furthermore, continuous tracking of other physiological parameters (e.g., glucose level), life style, diet, sleep, mood, behavior in the social media, and environmental conditions can be useful to define the entire picture of the patient’s condition and select the treatment regimen more precisely [33]. A continuous monitoring of the critical health parameters would mean a constant generation of a huge amount of data that has to be securely and efficiently processed and stored in a suitable online platform (e.g., cloud). The controlled access to the data would be granted only to the authorized parties.

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Automated decision-making based on machine learning algorithms and mathematical models derived from the known data sets can be introduced to process the input data and make the treatment related decisions about, for instance, the adjustment of the dose and medicine(s). Furthermore, different apps/chats based on the artificial intelligence can be used to give patients the possibility to communicate about the data generated from their wearable sensors in between the visits to the healthcare professionals. There is an increasing pressure to decrease the cost of the public health-care system, and more discussion on the ethics of the automated decision-making processes is needed. This discussion should be initiated already now, before we reach the same level of technical implementation that has been reached with many of the social media platforms, such as Facebook. The first examples of data leakage and larger systematic misuse of personal data, as seen with Facebook and Cambridge Analytica, underpin the importance of a careful design of health big data solutions.

2.3 Where do we need the 3D printing? The fabrication of patient-tailored medicine based on the continuous monitoring of the different parameters would require flexible technological solutions, because the conventional manufacturing of oral dosage forms is rigid, and with a very limited possibility for on-demand accurate and reliable production of personalized doses. 3D printing can offer the solution to this challenge, because it enables fast and precise fabrication of the specific doses by, for example, a simple manipulation of the physical dimensions in the software of the printer: higher dose would require a bigger tablet and vice versa [34]. In addition, oral dosage forms in various shapes, with or without a modified release profile can be manufactured in a single step using the same 3D printer [16]. Furthermore, pharmacoprinting holds a potential for oral multi-drug dosage forms that can be beneficial to boost polypharmacy patients’ compliance to the medicine. The last, but not the least, a small footprint of the 3D printer and the possibility of remote control over the equipment give the prospect for on-demand printing in the point-of-care and emergency settings [16]. On-demand printing can be especially beneficial for drugs with limited shelf lives.

2.4 Patient perception and preferences of the medicine regarding shape, color, embossing, flavor, and acceptability of patient-designed medicine The versatility of 3D printing includes the possibility of customizing the physical appearance of the dosage form. Color, shape, size and embossing design can be tailored to the preference of each patient to improve compliance. Brieger et al. showed that color could play a dramatic role in patient perception of the efficacy of the anti-malaria therapy [35]. The study conducted in Nigeria showed that, without a proper education, yellow prepacks would be considered as more effective than the blue ones [35]. The reason for such finding is the patients’ association of the disease with yellow color, as yellow eyes is the

3D printing in oral drug delivery

common symptom of malaria. Color can give meaning and power to the medicine even if it is a placebo. Red would be associated with stimulative effect, while blue with antidepressive action [35]. In general, patients prefer colors that do not provoke negative associations and emotions. Therefore, black is considered to be the least favorable color for the dosage form due to its relation to grief and death [36]. In countries where counterfeiting of drugs is a common practice, colored medicine could be preferred, because white could be considered to be only chalk. Moreover, Goyanes et al. have studied the acceptability of different shapes of dosage forms by patients. It was found that torus together with the conventional geometries, such as disc and capsule, were the most preferred shapes among the respondents, because of ease in swallowing and picking [36]. Sphere was considered unfavorable due to difficulties in grabbing it. Well-thought shapes can add extra functionality to the medicine. For instance, heart-shaped tablet is well-received by patients, because the shape is selfexplanatory in the sense of medicine indication. Embossing design on the dosage forms can be useful to specify the time of administration: sun for morning intake and moon for evening intake [37]. Other embossing designs used in this study, such as written name or abbreviation of time, were found irrelevant. Though, it was expressed that nurses and health assistants could find it useful in order to minimize the medical errors occurring when the tablets are distributed to the patients. Furthermore, the flavor of the dosage form can be preselected before the 3D printing [38]. This can be especially relevant if the drug product is to be distributed in different countries, because culture differences regarding the perception and preferences of flavor and taste can be huge. The potential of 3D printing could allow selecting the appearance of patient-centric medicine by patients, and bring the concept of medicines being designed “for the patient by the patient” in light. The customized appearance of the dosage forms helps patients to recognize and differentiate their medicine more easily, and avoids incorrect administration of the medicine. Furthermore, the adherence to the treatment can be improved if patients perceive their medicine as appealing, and easy to swallow and handle. In addition, patients showed interest to design their own medicine. The researchers believe that the involvement of patients in the designing process could further improve their adherence to the therapy [37].

3. Introduction to additive manufacturing 3.1 Fused deposition modeling Fused deposition modeling (FDM) was originally developed by Scott Crump at Stratasys and is one of the most extensively used additive manufacturing techniques. The FDM manufacturing principle is based on the extrusion of thermoplastic materials and subsequent deposition of the semi-molten materials onto a build plate in a layer-by-layer

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fashion. The materials used in FDM 3D printing, are usually conveyed by two opposing rollers to the nozzle of the extruder print head. In the print head, the materials are heated by thermocouple units to a semi-molten state. The print head will subsequently trace the volume of the intended print geometry horizontally while the semi-molten materials are extruded from the nozzle. After completion of the horizontal layer, the stage lowers and the subsequent layer is deposited. These stage lowering and deposition steps are subsequently repeated until the 3D structure has been created (Fig. 8) [39]. For FDM 3D printing there exist a number of parameters that may be adjusted and controlled in order to optimize the print quality and the print time. These parameters are, among others, the layer height, the velocity of the printhead, the speed of the rolls and the infill density [40]. The infill density defines the amount of polymer that is filled into the geometry, and this parameter subsequently controls the overall porosity of the 3D printed geometry. The infill density can be adjusted from 0% to 100%, where a 0% infill density creates a completely hollow object and 100% infill density creates a solid object [41]. The printing temperature is highly dependent on the thermoplastic polymer and presence of plasticizers. This is due to the fact that the thermoplastic polymers can have different inherent viscosities. These viscosities will be lower in the presence of plasticizers

Fig. 8 Outline of the FDM 3D printing technique where thermoplastic material is heated in the printhead and deposited on a movable build platform. (Reprinted with permission from the American Chemical Society (ACS): Gross BC, Erkal JL, Lockwood SY, Chen C, Spence DM. Evaluation of 3D printing and its potential impact on biotechnology and the chemical sciences. Anal Chem 2014;86:3240–53. https://doi.org/10.1021/ac403397r.)

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and hence, the processing temperature can be lowered [42]. Polyvinyl alcohol (PVA) is one of the often used thermoplastics in FDM 3D printing [43]. PVA is a water-soluble polymer that swells upon contact with aqueous media. Depending on the PVA hydrolysis degree of the acetate groups, the melting point may range from 180 °C (partially hydrolyzed) to 228 °C (fully hydrolyzed). PVA has been used as a filament, where drugs have been loaded into the filament by impregnation or where the drugs have been incorporated into the filament by hot melt extrusion (HME). The impregnation procedure is a common method that incorporates drugs into the filament and poses a limited risk for degradation [44]. For instance, PVA filaments have been loaded with prednisolone via the impregnation technique in a saturated methanol solution for 24 h. This impregnation technique achieved a loading of 1.9% w/w of prednisolone with a concomitant minor degradation of the drug. The impregnated filaments had to be subsequently dried in an oven. This impregnated filament could then be utilized to print tablets. Due to the high temperature of the printing process and the chosen loading method, prednisolone was found to be in the amorphous state in the PVA polymer [34]. Similar impregnation methods have been performed on 4-aminosalicylic acid (4-ASA), 5-aminosalicylic acid (5-ASA, mesalazine) and fluorescein with drug loading % between 0.06% and 0.29% w/w. It was also observed that especially 4-ASA experienced degradation due to contact with the heated extruder of the printer. All the incorporate drugs in the printed tablets displayed modified-release profiles, which could be modified in the case of 4-ASA to last more than 7 h. It was also observed that a higher infill density drastically extended the release of fluorescein (fluorescein dissolution completed after 20 h for 90% infill density, 15 h for 50% infill density and 6 h for 10% infill density) [18, 45]. The passive diffusion that needs to occur in the impregnation method necessitates the use of highly concentrated solutions of drugs in order to incorporate limited amounts of drug into the thermoplastic filament. This renders the impregnation method as being both a time consuming, waste generating and expensive process. Furthermore, the utilized drugs should also be stable when dissolved in the solvent, sufficiently thermostable and the solvent should not damage or dissolve the mechanical or physical integrity of the thermoplastic filament. Therefore, it has appeared to be useful to develop a manufacturing approach that utilizes a combination of HME and FDM 3D printing [44]. Theophylline-loaded delivery systems based on HME and FDM 3D printing of theophylline, Eudragit® and plasticizer constituents have achieved approximately 50% drug loading [46]. Besides having the possibility to either impregnate or hot melt extrude the drug with thermoplastic filament, it is also possible to 3D print hollow capsule shells with compartments for deposition of the drugs. The compartments of the drug delivery devices can be fabricated using different materials and/or thicknesses, thus modifying the drug release profile from each of the compartments [47]. The drug filling of the compartments can also be performed during FDM 3D printing, and as they are thus sealed, the compartments can be filled with extrudates, powders or even liquids [27, 28, 48].

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Furthermore, a change of drug or formulation composition can be performed using this FDM 3D printing filling approach without affecting the printing process, which is not the case if the API is incorporated into the filament prior to printing using either impregnation or HME. Furthermore, having the drug loading step spatially decoupled from the FDM 3D printing step limits the API exposure to high temperatures [28]. Finally, it is also a desirable attribute of this method that two or more drugs can be co-formulated even if they are chemically incompatible, since they can be separated in different compartments [27]. The FDM 3D printing method is hence a process that offers an alternative to conventional preparation methods. The release profiles of the incorporated drugs can be adjusted freely by modifying the geometrical shape of the drug delivery systems [20].

3.2 Semi-solid extrusion The semi-solid extrusion printing method functions via extrusion of a viscous semiliquid material from a syringe and hence resembles the FDM 3D printing, although often performed at a lower temperature (Fig. 9). The semi-solid extrusion process can be performed at ambient temperature [49]. The rheological properties of the extrusion materials are considered to be the major parameters that have to be controlled to obtain reproducible DDS. These rheological properties are highly dependent on the solid material loading amount, and type and amount of the additives [42]. This semi-solid extrusion method has been utilized to manufacture a complex multidrug tablet, where three drugs were capable of being released either by diffusion or by osmosis [49]. The osmotic release profiles were obtained by making a semi-permeable membrane consisting of cellulose acetate and PEG 6000. When the dosage form was

Fig. 9 Outline of the semi-solid extrusion 3D printing technique: (A) a photograph and (B) schematic illustration of the 3D printer. (Reprinted with permission from Elsevier: Khaled SA, Burley JC, Alexander MR, Roberts CJ. Desktop 3D printing of controlled release pharmaceutical bilayer tablets. Int J Pharm 2014;461:105–11. https://doi.org/10.1016/J.IJPHARM.2013.11.021.)

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exposed to contact with aqueous media, the hydrophilic constituent dissolved and the pores were subsequently formed. These pores were capable of controlling the diffusion of water through the osmotic unit. The connecting layer between the compartments was composed of polyvinyl pyrrolidone (PVP) K30, mannitol and surperdisintegrants, thus providing a fast disintegration of the dosage form. The diffusional release of nifedipine and glipizide was obtained using a hydrophilic hydroxypropyl methylcellulose-based matrix that formed a gel. Prior to dissolution testing the dosage form had to be dried for 24 h [49]. This technique has also been utilized to manufacture a polypill containing five compartmentalized drugs with both sustained and immediate release profiles [24]. A downside to this technique is the use of solvents, which may necessitate the implementation of expensive analytical methods for quantifying their removal during the drying step, and the drying step may be a slow procedure. Furthermore, the slow solidification of the dosage form and shrinking of the matrix gel may affect the final properties of the 3D printed dosage form [44].

3.3 Powder-bed printing The powder-bed 3D printing method is utilizing a combination of a powder-bed and a binder ink to build a solid structure layer by layer. The powder-bed 3D printing dosage forms typically consist of layers with a height of 200 μm, with powder particle sizes between 50 and 100 μm [50]. These particles and particle layers are bound together with the printed ink to generate a 3D model. The initiation of the process starts by distributing an even layer of powder on a support stage, usually by a roller application method. Subsequently, the inkjet printer head adds droplets of liquid binding material onto the evenly distributed powder layer at the intended areas where solidification is required. When this addition of liquid binding material has completed, the stage lowers with a predetermined step size, and a subsequent powder layer is evenly distributed on top of the initial layer and again selectively sprayed with the printed binding material. This sequence is iterated until the desired geometrical shape of the DDS is generated. The DDS is then usually heated to intensify the binding strength of the powders where liquid binding material has been deposited. The DDS can subsequently be removed from the powder-bed wherein the unbound powder acts as a support material during the manufacturing process (Fig. 10) [39]. Besides the properties of the powder and ink and the shape of the DDS, multiple process parameters may also affect the physicochemical properties of the system. The layer spacing and thickness between the printed lines has to be optimized to obtain an adequate adhesion between the individual layers. The x–y axes speed and flow rate of the binder ink have a direct influence on the amount of binder that is deposited per unit line length. As an example, a decrease of the axis speed while the flow rate is increased produces tablets that have a lower friability and a greater strength. However, depositing larger

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Fig. 10 Outline of the powder-bed 3D printing technique. (Reprinted with permission from ACS: Gross BC, Erkal JL, Lockwood SY, Chen C, Spence DM. Evaluation of 3D printing and iIts potential impact on biotechnology and the chemical sciences. Anal Chem 2014;86:3240–53. https://doi.org/10.1021/ ac403397r.)

quantities of binder may yield a DDS with poor surface resolution. The powder-bed system can incorporate complex internal attributes, such as a porosity gradient, multiple material regions, and tortuous channels. Highly complex release profiles can be obtained with this technique, such as multiple-release patterns stemming from one single dosage form. Such intricate release profiles can be easily produced using this technology. Conventional tablet compression methods may be capable of providing only a limited degree of spatial control (e.g., bilayer tablet production), but the conventional tablet compression methods are incapable of reproducing the convoluted geometry available with powder-bed printing [15].

4. Regulatory challenges 4.1 The first 3D printed drug product on the market 3D printing possesses the unique possibility for on-demand production of personalized doses to fulfill the need of each individual. Patient-tailored medicine implies, in most cases, a continuous adjustment of the dose based on the response received from a continuous monitoring of different parameters, as discussed above. For instance, by measuring the level of glucose, the dose of insulin for injection can be adjusted. However, dose adjustment is not as simple for medicines available as oral solid dosage forms. Oral solid

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dosage forms are registered in the discrete doses, and the provision of the correct dose is limited to the possible combinations of the available strength or half of the strength in home conditions. There are several factors contributing to the fact that the large-scale production of pharmaceuticals is only done based on the fixed dose assumption. The regulatory framework for manufacturing of products with a flexible dose, except for extemporaneously prepared drug products, is not straightforward. The first and only FDA approved 3D printed oral solid dosage forms, Spritam®, containing antiepilepsy drug levetiracetam, is only available as four discrete doses, despite the unique personalization possibilities of 3D printing.

4.2 Requirements for raw materials, printers and manufacturing procedures Tremendous amount of research work to explore the possibility for manufacturing of oral solid dosage forms has been done with raw materials that are not suitable for oral consumption [51]. Currently, none of the commercially available filament materials used, for example, in FDM are suitable to be orally administered. Only a limited amount of the pharmaceutically approved excipients and their combinations can be successfully used for the currently available 3D printers. None of the commercially available printers is directly intended to be used for pharmaceutical purposes [52]. Furthermore, before the printer can be used to fabricate drug products, the printer and the printing environment should be certified according to the principles of good manufacturing practice (GMP). There are commercially available printers that can be used for the commercial production of highresolution parts for cars and airplanes. Therefore, the challenge is to motivate the companies to design printers intended for the pharmaceutical use. One promising direction is the use of extrusion-based feeding for the 3D printer [53]. In other words, instead of using pre-made filaments to be loaded to the 3D printer, raw material in powder/pellet form would be directly fed to the printer. However, this situation could also provoke the misuse of printing technology with supply of falsified drug products, for instance, through online pharmacies. As personalized dosage units could look different, depending on the dose and the desired design, the genuineness of the supplied drug product has to be verified. Introduction of track-and-trace elements onto/into 3D printed dosage units could be an option [54]. The concept of cryptopharmaceuticals has also been introduced to allow for increased traceability of medication at a single dosage unit level [55]. The on-demand production of flexible doses to meet the needs of individual patients would require fast and inexpensive quality assurance systems to verify that a correct drug in a correct dose is printed. Non-destructive quality control tools, such as spectroscopic, colorimetric and smart mobile devices would come into play [56]. Sensitive miscroscales can also be used to verify the printed dose. Routine validation procedures would be needed to ensure the correct performance of the printer. Regular maintenance of the

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production line should be performed. A duplicate of the printer in use would be needed to ensure a nonstop supply of the medicine even if the printer is malfunctioning [57]. It is assumed that a basic setup of the printer with interchangeable parts could be used to print different drug products. In this scenario, cross-contamination has to be avoided. The proper cleaning procedures have to be introduced. If solvents are involved in the printing process, the procedure for their proper removal should be specified, together with a proper overall waste management plan. In case of powder-bed printing, generation of toxic and explosive dust could represent health and occupational hazards [58]. The physical, chemical and microbiological stability of the ingredients in the dosage forms and the entire dosage unit should be verified for the intended storage time according to the existing regulatory guidelines. However, in the case of on-demand personalized production, the expected shelflife of a product can be significantly shorter than for traditionally manufactured pharmaceutical products.

4.3 Supply chain The introduction of 3D printing as a flexible manufacturing technique for on-demand production of personalized medicine could bring a change in the drug distribution chain [33]. Particularly, the production of flexible dosage units could happen in the pharmacy or hospital (Scenario A) [38], patient’s home (Scenario B) or pharmaceutical industry (Scenario C) (Fig. 11). Compounding pharmacies and hospitals (Scenario A) are logical places for the flexible production of pharmaceuticals based on the continuous monitoring of the critical parameters, because both settings have experience with compounding. Additionally, patients Current

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Fig. 11 Representation of the current distribution chain of the drug products and three possible scenarios for future logistics of on-demand produced personalized medicine. Red line indicates the place for manufacturing of flexible dosage units. (Reprinted with permission from Taylor & Francis Online: Lind J, Ka€lvemark Sporrong S, Kaae S, Rantanen J, Genina N. Social aspects in additive manufacturing of pharmaceutical products. Expert Opin Drug Deliv 2017;14:927–36. https://doi.org/10. 1080/17425247.2017.1266336.)

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are used to receive drug products from the pharmacy. They trust the quality of medicines and the related advice. The challenges associated with 3D printing of personalized medicine in the pharmacy or hospital are related to the fact that new knowledge would be required from the personnel. Pharmacies are usually limited in space, and crosscontamination of the materials between different formulations is possible. With the provision of affordable printers, flexible production of personalized medicine could potentially happen at home (Scenario B), which would be a practical scenario for many patients. However, it could be very dangerous to deal with the raw materials at home, even if they are distributed as the intermediate product, such as the drugcontaining filament for FDM. Furthermore, the question remains on how the validation and maintenance of the printer would be performed. Definitely, patients have to be educated before the printer could be located at home, even if it would be operated remotely. The pharmaceutical industry has naturally the facilities for the production of medicines, and could potentially be able to produce personalized doses on-demand, as they also possess the required knowledge and level of security. However, there are major challenges when considering a direct contact with a patient and solving the questions related to private data. On-demand pharmacoprinting as a new approach would require economic evaluation of its benefits. It is still unclear whether 3D printing of personalized medicine would be a valuable alternative to the existing manufacturing methods of drug products. One of the key challenges is the liability question [58]. Additionally, the role of the insurance companies and/or public insurance system should be debated in any of the possible scenarios. Additionally, a wide availability of personalized medicine could be accompanied with unintended consequences for patients [59]. For instance, not all patients are capable and interested in everyday measurements of different health parameters. Furthermore, patients could deny accepting a pharmacoprinted drug, because of its different appearance from the traditional solid dosage form.

4.4 Data management from the regulatory perspective On-demand production of personalized medicine requires a new type of thinking in the regulation of medicine. It is a huge change to move from regulation of batch-wise production of medicinal products with a precise (e.g., 5 mg) content toward continuously manufactured personalized doses that can range, for example, from 2 to 20 mg. As pointed out earlier, tracing these products requires new solutions and ways of handling the single product and production related data. The health-related big data will be a key for the successful implementation of personalized medicine. One of the key challenges will be the storage of the personal data and the final database structure. The data need to be accessible for multiple actors in the health care sector, but at the same time, highly protected against cyberattacks. Pharmaceutical

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cybersecurity will be an increasingly important area in the future and should be also built into the future regulatory paradigm around personalized medicine. Another important question is the physical location of the data, because the required size of the computer clusters and supercomputers is remarkable. One trend in the data science area has been to find the cheapest solution, but in the case of health data, special attention should be paid on ensuring the cybersecurity, by placing the supercomputer into an environment where the infrastructure is capable of ensuring the protection of the data. Another challenge is the regulation of automated decision-making processes related to critical decisions on treatment strategies. This automated decision-making is based on algorithms using the patient specific data. The related algorithms typically involve not only artificial intelligence (AI) and other less traditional mathematics, but also deep learning algorithms and methods that would require mathematical skills far outside the typical pharmaceutical curriculum and a classical regulatory skill set. These algorithms will be a part of mobile medical applications, where the first regulatory documents have been introduced recently (e.g., FDA guidance “Policy for Device Software Functions and Mobile Medical Applications”). Solving this requires a new thinking in the field of education of pharmaceutical experts. Data rich pharmaceutical supply chain will involve storage of sensitive personal data, and the protection of that is of major importance. The final data analytical solutions that will be used to personalize the overall treatment strategy should be compliant with initiatives such as the European General Data Protection Regulation (GDPR).

5. Future perspective 5.1 Robotic devices Additive manufacturing provides the possibilities that the conventional manufacturing methods lack. 3D printing could be used to produce oral robotic devices that would release the drug triggered by internal or external stimulus. Oral self-oriented devices for injections of the medicine in the stomach tissue, activated by internal stimulus, i.e., GI liquid, can be applied for drugs whose bioavailability would benefit from such a delivery route [6]. Furthermore, other parts of the GI tract can be targeted for the delivery of the pharmaceuticals directly to the tissue, without a direct contact with the harsh conditions of the GI tract. For instance, injection of the active ingredients into the colonic tissue after oral uptake could be triggered by the colonic bacteria. Oral origami robots can be the future of targeted drug therapy [60]. The robots are folded many times to allow their encapsulation into objects such as ingestible capsules to ease the swallowing of the automatic machine. These devices, manipulated by the external magnetic field, could be used to deliver the drug exactly to the desired location in the GI tract, independently of the transit time. Furthermore, the unique design of the medical devices (e.g., microcontainers) can allow

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their prolonged deposition at the desired location/membrane in the GI tract and the unidirectional release of the active compound, protecting the drug from harsh GI conditions [61].

5.2 Digital technology and Internet of Things: Fabrication, identification, anti-counterfeiting The widespread use of digital technologies also affects the healthcare sector. A smartphone is becoming an indispensable tool to help patients to manage their disease conditions. Different apps are made available for self-monitoring of the critical disease parameters. This monitoring is possible through Internet of things (IoT), such as wearable pulse rate meters, lenses with the ability of measuring intraocular pressure, and glucose meters that continuously supply data to the designed apps without the need for human intervention. Potentially, these data from the apps can be used to find out and/or adjust the correct dose of the correct oral medicine. Here, additive manufacturing comes into play with its flexibility in the production of the desired dose on-demand. However, such flexible fabrication of the doses would require that a patient could identify his/her medicine correctly, especially if a polypharmacy patient, who has to consume multiple drugs at the same time, is in question. A digital solution could be to mark the dosage forms with a scannable quick response (QR) code [54] or even inkjet print the drug in the pattern of a QR code [62]. The latter QR pattern would contain simultaneously the drug and information relevant to the end-user, and more. The patient would have to scan the QR encoded dosage form with a normal smartphone before consumption for identification purposes. This could help patients to ensure that the correct medicine, in the right dose and at the right time of the day would be taken. Furthermore, QR encoding of each dosage unit would allow tracing the distribution chain of the medicine from the manufacturing unit to the end-user. This could help in the fight against counterfeiting of medicines that possess a serious threat to global health nowadays. To reduce a significant spread of the counterfeit medicine, in February 2019 the European Commission has issued the Falsified Medicine Directive, according to which the authenticity of each medicinal product manufactured and distributed in Europe has to be verified [63]. This item-level serialization poses a challenge for each player in the supply chain as the current pharmaceutical products do not straightforwardly comply with this requirement. Therefore, the fabrication of the oral dosage units in the pattern of a QR code could solve this problem. Furthermore, it could help to avoid drug shortage that can leave a patient without the required drug product for a while. In addition, the resale of the tainted drug products can be circumvented if each drug product is tracked-n-traced in the systems that different players of the supply chain have access to. Moreover, a constant digital tracing of the consumable dosage forms together with the continuous monitoring of the critical health parameters can be done by both patients and healthcare professionals. This could

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be especially important when the administration of drugs with potential misuse is in question. The early detection of drug abuse could avoid the fatal outcome of the therapy. The global digitalization with transfer and storage of private data makes the public key infrastructure vulnerable for cyberattacks. To minimize the threat against unauthorized access to the private user data collection with subsequent falsification of the information, a patient-unique blockchain of health data could potentially be used [55]. With such a solution, it would be very difficult to mess up the data.

5.3 Personalized medicine for veterinary purposes Esthetic appearance of things can be crucial in humans’ desire to possess, purchase and/or consume. Food companies use different shapes and colors on their products for marketing purposes. An increased variety in the design of the consumables is also seen with the dry food products intended for pets. For instance, dry formulations in the shape of bones, balls, and diamonds are available for dogs; fish-, mouse-, and flower-shaped are available for cats; and brightly colored flakes in the shape of heart, banana and grain are available for parrots. It gives an impression that the investment in personalized design of medicine for veterinary purposes would be a natural step for pharma business. Furthermore, a continuous monitoring of critical health parameters via IoT is possible with animals too. It means that a correct dose of a correct medicine can be provided to pets. A number of skin and coat conditions can be better treated by analyzing the images regularly taken with a smartphone camera, by using an appropriate app.

6. Concluding remarks The pharmaceutical field is facing a historical challenge: genetic information and wearable point-of-care technologies can be used for precisely defining the dose of medicine for each individual patient. However, the existing manufacturing solutions and supply chain models are the rate-limiting step for the implementation of new product design principles fitting into this picture. By the integration of data sciences and new process engineering principles, innovative medicinal products can be introduced. At the same time, it is important that we simultaneously develop the regulatory framework for personalized products, as well as continue the public debate on the use of personal data in this context.

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[41] Mohanty S, Larsen LB, Trifol J, Szabo P, Burri HVR, Canali C, Dufva M, Emneus J, Wolff A. Fabrication of scalable and structured tissue engineering scaffolds using water dissolvable sacrificial 3D printed moulds. Mater Sci Eng C 2015;55:569–78. https://doi.org/10.1016/J.MSEC. 2015.06.002. [42] Aho J, Boetker JP, Baldursdottir S, Rantanen J. Rheology as a tool for evaluation of melt processability of innovative dosage forms. Int J Pharm 2015;494:623–42. https://doi.org/10.1016/j.ijpharm.2015.02.009. [43] Kollamaram G, Croker DM, Walker GM, Goyanes A, Basit AW, Gaisford S. Low temperature fused deposition modeling (FDM) 3D printing of thermolabile drugs. Int J Pharm 2018;545:144–52. https:// doi.org/10.1016/J.IJPHARM.2018.04.055. [44] Goole J, Amighi K. 3D printing in pharmaceutics: a new tool for designing customized drug delivery systems. Int J Pharm 2016;499:376–94. https://doi.org/10.1016/J.IJPHARM.2015.12.071. [45] Goyanes A, Buanz ABM, Hatton GB, Gaisford S, Basit AW. 3D printing of modified-release aminosalicylate (4-ASA and 5-ASA) tablets. Eur J Pharm Biopharm 2015;89:157–62. https://doi.org/ 10.1016/j.ejpb.2014.12.003. [46] Pietrzak K, Isreb A, Alhnan MA. A flexible-dose dispenser for immediate and extended release 3D printed tablets. Eur J Pharm Biopharm 2015;96:380–7. https://doi.org/10.1016/j.ejpb. 2015.07.027. [47] Maroni A, Melocchi A, Parietti F, Foppoli A, Zema L, Gazzaniga A. 3D printed multi-compartment capsular devices for two-pulse oral drug delivery. J Control Release 2017;268:10–8. https://doi.org/ 10.1016/J.JCONREL.2017.10.008. [48] Okwuosa TC, Soares C, Gollwitzer V, Habashy R, Timmins P, Alhnan MA. On demand manufacturing of patient-specific liquid capsules via co-ordinated 3D printing and liquid dispensing. Eur J Pharm Sci 2018;118:134–43. https://doi.org/10.1016/j.ejps.2018.03.010. [49] Khaled SA, Burley JC, Alexander MR, Yang J, Roberts CJ. 3D printing of tablets containing multiple drugs with defined release profiles. Int J Pharm 2015;494:643–50. https://doi.org/10.1016/j. ijpharm.2015.07.067. [50] Pfister A, Landers R, Laib A, H€ ubner U, Schmelzeisen R, M€ ulhaupt R. Biofunctional rapid prototyping for tissue-engineering applications: 3D bioplotting versus 3D printing. J Polym Sci Part A Polym Chem 2004;42:624–38. https://doi.org/10.1002/pola.10807. [51] Jamro´z W, Szafraniec J, Kurek M, Jachowicz R. 3D printing in pharmaceutical and medical applications—recent achievements and challenges. Pharm Res 2018;35:176. https://doi.org/ 10.1007/s11095-018-2454-x. [52] Aho J, Bøtker JP, Genina N, Edinger M, Arnfast L, Rantanen J. Roadmap to 3D-printed oral pharmaceutical dosage forms: feedstock filament properties and characterization for fused deposition modeling. J Pharm Sci 2019;108:26–35. https://doi.org/10.1016/J.XPHS.2018.11.012. [53] Welsh NR, Karl Malcolm R, Devlin B, Boyd P. Dapivirine-releasing vaginal rings produced by plastic freeforming additive manufacturing. Int J Pharm 2019;118725. https://doi.org/10.1016/j. ijpharm.2019.118725. [54] Trenfield SJ, Xian Tan H, Awad A, Buanz A, Gaisford S, Basit AW, Goyanes A. Track-and-trace: novel anti-counterfeit measures for 3D printed personalized drug products using smart material inks. Int J Pharm 2019;567:118443. https://doi.org/10.1016/J.IJPHARM.2019.06.034. [55] Nørfeldt L, Bøtker J, Edinger M, Genina N, Rantanen J. Cryptopharmaceuticals: increasing the safety of medication by a blockchain of pharmaceutical products. J Pharm Sci 2019. https://doi.org/10.1016/ j.xphs.2019.04.025. [56] Edinger M, Jacobsen J, Bar-Shalom D, Rantanen J, Genina N. Analytical aspects of printed oral dosage forms. Int J Pharm 2018;553:97–108. https://doi.org/10.1016/J.IJPHARM.2018.10.030. € [57] Oblom H, Sj€ oholm E, Rautamo M, Sandler N. Towards printed pediatric medicines in hospital pharmacies: comparison of 2D and 3D-printed orodispersible warfarin films with conventional oral powders in unit dose sachets. Pharmaceutics 2019;11:334. https://doi.org/10.3390/ pharmaceutics11070334. [58] Alhnan MA, Okwuosa TC, Sadia M, Wan KW, Ahmed W, Arafat B. Emergence of 3D printed dosage forms: opportunities and challenges. Pharm Res 2016;33:1817–32. https://doi.org/10.1007/s11095016-1933-1.

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[59] Kaae S, Lind JLM, Genina N, Sporrong SK. Unintended consequences for patients of future personalized pharmacoprinting. Int J Clin Pharmacol 2018;40. https://doi.org/10.1007/s11096-018-0596-x. [60] MIT. Ingestible origami robot. MIT News; 2016 [WWW document], http://news.mit.edu/2016/ ingestible-origami-robot-0512#news-video-block. (Accessed 14 August 2019). [61] Nielsen LH, Nagstrup J, Gordon S, Keller SS, Østergaard J, Rades T, M€ ullertz A, Boisen A. pH-triggered drug release from biodegradable microwells for oral drug delivery. Biomed Microdevices 2015;17:55. https://doi.org/10.1007/s10544-015-9958-5. [62] Edinger M, Bar-Shalom D, Sandler N, Rantanen J, Genina N. QR encoded smart oral dosage forms by inkjet printing. Int J Pharm 2017. https://doi.org/10.1016/j.ijpharm.2017.11.052. [63] European Commision. Falsified medicines. Public Health; 2019 [WWW document], https://ec. europa.eu/health/human-use/falsified_medicines_en. (Accessed 21 May 2019).

Further reading Khaled SA, Burley JC, Alexander MR, Roberts CJ. Desktop 3D printing of controlled release pharmaceutical bilayer tablets. Int J Pharm 2014;461:105–11. https://doi.org/10.1016/J.IJPHARM.2013.11.021. Servier Medical Art [WWW document], https://smart.servier.com/smart_image/complete-digestiveapparatus-3/; 2019. (Accessed 29 October 2019).

CHAPTER 13

3D intestinal models towards a more realistic permeability screening Andreia Almeidaa,b,*, Cláudia Azevedoa,b,*, Maria Helena Macedoa,b,*, Bruno Sarmentoa,c a

i3S—Instituto de Investigac¸a˜o e Inovac¸a˜o em Sau´de, Universidade do Porto, Porto, Portugal Instituto de Ci^encias Biomedicas Abel Salazar, Universidade do Porto, Porto, Portugal c CESPU, Instituto de Investigac¸a˜o e Formac¸a˜o Avanc¸ada em Ci^encias e Tecnologias da Sau´de, Gandra, Portugal b

1. Introduction Three-dimensional models, although being a relatively new topic, have been explored in the past years, since they can bring new insights and valuable knowledge when studying a specific organ. Regarding drug development, and specifically oral drug delivery, the need for an intestinal model that can better mimic the human small intestine is unquestionable. Truth is, models that are currently used to test the permeability’s potential of new formulations in vitro are very limited. In 2006 it was reported that if there was a 10% improvement in the preclinical screening methods used, this could reduce the costs of drug development by 100 million dollars per drug approved to reach the market [1, 2]. In fact, drug testing in 2D systems provides information that can often be misleading, since the way cells are grown affects how they respond pharmacologically, meaning that 3D multicellular models can help bridge the gap between in vitro and animal testing [3]. The “gold-standard” of intestinal in vitro models is the Caco-2 model (Fig. 1A), which is comprised of only one cell type—Caco-2, a cell line derived from human colorectal adenocarcinoma. It is described that after 21 days in culture, these cells can differentiate into enterocyte-like cells, the major population of the intestinal epithelium [4]. However, this model possesses several limitations. Besides presenting different levels of drug transporters and metabolizing enzymes from what is observed in vivo, these cells form a monolayer with tighter tight junctions (TJ), that can lead to erroneous results, especially when testing the permeability of compounds that are transported paracellularly [5]. Researchers have been trying to improve the Caco-2 model, mainly by cultivating other cell lines that could improve the robustness of the model. Currently, the triple co-culture model is very used when testing intestinal permeability and consists of Caco-2 cells co-cultured with HT29-MTX and Raji B cells (Fig. 1B) [6]. HT29-MTX is a modified cell line from the HT29 cell line, with origin in human colorectal adenocarcinoma, treated in a medium * Equal contribution. Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00003-X

© 2020 Elsevier Inc. All rights reserved.

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Transwell® insert Basolateral chamber

Caco-2 Clone

Apical chamber

HT29-MTX M-cell

Permeable membrane

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Fig. 1 2D intestinal models that are currently used to perform permeability studies. (A) Representation of the “gold-standard” Caco-2 model. (B) Triple model comprising Caco-2, HT29-MTX and Raji B cells.

containing 10
6 M methotrexate (MTX) to become mucus-producing cells [7]. Another feature of these cells is that they possess looser TJ and, when co-cultured with Caco-2, promote a decrease in the transepithelial electrical resistance (TEER) of the monolayer, becoming more similar to what is observed in vivo being, nevertheless, higher than what is observed in the human small intestine. The production of mucus by these cells is also very important, since the small intestinal epithelium is covered by a mucus layer that confers protection but can also act as a barrier to the permeability of drugs [8]. The triple model is normally completed by the addition of Raji B cells on the basolateral side on the 14th day of culture, since these cells promote the differentiation of a fraction of Caco-2 cells into M-cells, which are present in the Peyer’s patches and are known to initiate mucosal immunity response, being specialized in transepithelial transport [6, 9]. Researchers have been trying to find alternatives to the Caco-2 cell line because of the aforementioned limitations. Different cell lines have been tested, like Madin-Darby canine kidney (MDCK), Lewis lung carcinoma-porcine kidney (LLC-PK), TC-7 (aCaco-2 sub clone), T-84 derived from lung metastasis of colon carcinoma, and IEC-18 and 2/4/A1, both intestinal rat cell lines [5, 10–12]. However, because of their origin or their own limitations, none has proven to be a good substitute for the Caco-2 cell line. In fact, any cell cultured on a flat surface, like the Transwell inserts membrane that are normally used for permeability assays, will never be able to fully replicate the in vivo situation. Cells, in vivo, are not flat, but in a three-dimensional configuration [1]. The components of the extracellular matrix (ECM) and the cell-to-cell and cell-to-matrix interactions are very important in cell behavior, but in 2D cultures these properties are lost [13]. In the case of the small intestine, its composition, functions and conditions make it a complex organ, responsible for the digestion of food and its absorption, as well as the absorption of drugs that are orally administered [14]. The human small intestine is divided into four layers: the mucosa, the submucosa, the muscularis propria and the serosa [15]. The mucosa is the most complex layer, comprised by the epithelium, the lamina propria and the muscularis mucosae, and is where the absorptive function takes place. In the

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epithelium, two populations are present. The absorptive enterocytes, that account for more than 80% of the population, are mainly responsible for the absorption of compounds. The other population present in the epithelium are the secretory cells, which include the goblet cells that produce the mucus layer that protects the epithelium from external factors, having an impact in the absorption of compounds; the Paneth Cells that secrete proteins that have a role in intestinal immunity; enteroendocrine cells that are responsible for secreting different hormones; and M cells that are specialized in transepithelial transport, as it has already been stated [15]. The intestinal lamina propria forms villi structures, enhancing the absorption area [16]. The composition of the lamina propria and the mechanical properties influence the behavior of the epithelial cells. Besides, cells that reside in the intestinal lamina propria, like fibroblasts, myofibroblasts, and mesenchymal stem cells can also influence the behavior of the epithelium [17–19]. As so, it is important to replicate the 3D environment in order to have a model that can better replicate the in vivo situation. Some work in this field has been developed in the last years, with promising results. There is a variety of 3D intestinal models, from the ones that are based on scaffolds or hydrogels, to decellularized tissue models, to more complex gut-on-a-chip models and organoids. A review of these models will be provided in the next sections.

2. 3D intestinal models 2.1 Multilayered models Since the 3D support is so important for an accurate mimicry of the organ physiology, several researchers have been using 3D matrices to develop new intestinal models. These matrices are normally constituted by ECM components, being collagen the most widely used [18, 20–23]. The choice for collagen is justified by its low cost, the easiness in processing and the flexibility for live cell manipulation. The tunability of properties like pore size and stiffness is also a great advantage of this material [3]. Other materials, like alginate, Matrigel, synthetic polymers and decellularized ECM are also used [8, 17, 24–29]. In fact, the 3D environment that surrounds cells in vivo is extremely important in cell functioning. The mechanical properties of the substrate are dependent on its composition, architecture and degree of crosslinking, and these factors dictate how mechanical forces are transmitted to cells [30]. Cell adhesion and migration is also controlled by matrix composition, being that cells might be able to remodel it [30]. As so, it is important to take this into account when choosing what material to use in order to mimic an organ. Besides, composition of the material, its mechanical characteristics and response to cells should be addressed and understood in order to replicate as much as possible the in vivo environment.

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While some authors do not give much importance to the architecture of the scaffold, using it merely as a 3D support that is more natural to the cells [17, 20], others build scaffolds with the intestinal villi architecture, claiming it has an effect on cell behavior and permeability of compounds [8, 18, 23]. There are several studies using scaffolds with villi architecture and using different techniques to obtain these structures (Fig. 2). A study by Sung et al. [8] describes a simple method to fabricate natural and synthetic hydrogels into 3D geometries with high aspect ratio and curvature. Authors were able to obtain villi structures mimicking the density and size of human intestinal villi, which is complex, without damaging them, combining laser ablation and sacrificial molding techniques that minimize the stress associated from separating the mold from the hydrogel

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Fig. 2 Micromolding a collagen scaffold on a porous membrane. (A) Schematic of the process used to generate a poly(dimethylsiloxane) (PDMS) stamp to micromold the collagen. (B) The dimensions of the microstructures in the stamp, which replicate the features formed in the collagen. The units for the numbers are in mm. (C) A top view of the stamp is shown using brightfieldmicroscopy. (D) A side view of the stamp is shown using electron microscopy. (E) Schematic of micromolded collagen in the modified insert. (F) Top views of the shaped collagen array acquired by brightfield microscopy. (G) A close-up view. Reprinted with permission from Wang Y, et al., A microengineered collagen scaffold for generating a polarized crypt-villus architecture of human small intestinal epithelium. Biomaterials 2017;128:44–55.

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structure [8]. After obtaining the structure, Caco-2 cells were seeded and cultured for three weeks and were able to cover the structures [8]. Later, Yi and colleagues [21] compared the absorptive and metabolic properties of Caco-2 cells cultured on the collagen villi scaffold from the previous work [8] with those in a 2D monolayer. The results obtained were quite promising, showing that the activity of the intestinal epithelial differentiation marker alkaline phosphatase and the metabolic enzyme P450 3A4 were increased. On the other hand, the activity of dipeptidase was lower than in the 2D monolayer control. The authors also tested the permeability of different compounds and observed a higher permeability in the 3D model, which correlated better to the in vivo situation [21]. Yu et al. [18] also developed a villi structure replicating the shape and size of human small intestine villi using the method described by Sung et al. [8] and compared with 2D models regarding paracellular drug absorption and cell growth. In this case, the authors only cultured the cells on the scaffolds for 14 days to avoid degradation of the villi structure and multilayer formation. The permeability of Atenolol, a slowly absorbed compound was higher than in the 2D model and more similar to permeability coefficients for perfused human intestine obtained by Lennernas et al. [31]. They were also able to observe that cell differentiation on 3D villi scaffolds varied along the villous length, where cells were more polarized and columnar at the top and less differentiated near the villous base [18]. Using villi structures has other advantages like the fact that it is easier to subject the cells to gradients of growth factors that can promote the creation of a stem/progenitor cell zone and cell migration along the crypt-villus axis [23]. This easiness to apply gradients to cells was observed by Wang and colleagues [23], who tried to recreate the intestinal microenvironment to understand the mechanisms behind the differentiation of cells in this organ. Using a collagen scaffold that was micropatterned using stamps in order to mimic the villi and the crypts of the small intestine, it was observed that cells plated on this scaffold were guided to form a crypt-villus architecture. Furthermore, applying a gradient of growth factors and DAPT (gamma secretase inhibitor) to the culture, the cells became polarized, promoting the formation of a progenitor-cell zone in the crypt and differentiated cells in the villus axis [23]. As it can be observed from the previous study, although most authors focus on the villus, it is clear that the crypts also play an important role regarding cellular differentiation in the small intestine. Crypt influence was studied by Wang et al., [24] who fabricated poly(dimethylsiloxane) (PDMS) substrates with “holes” mimicking the intestinal crypts, using a mold obtained through lithography technique. The surface was coated with fibronectin, an ECM protein, and Caco-2 cells were seeded on the substrate. The authors observed that after 4–5 days in culture, cells were able to cover the entire surface, migrating from the bottom to the top of the well structures, whereas this spreading was slower comparing to flat surfaces. Nevertheless, the topography of the substrate affected cell metabolic activity and differentiation, showing an increase in mitochondrial activity and a decrease in alkaline phosphatase activity. These differences stress out the importance of topography, and it is

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important to notice that differences in cell behavior can lead to differences in results regarding permeability testing. Getting back to the villi, Costello et al. [25] focused on the development of a porous PLGA 3D scaffold, also with villi architecture. The scaffold supported the growth of epithelial Caco-2 and HT29-MTX cell lines. In this study, authors were also able to observe that mimicking the intestinal topography, cellular differentiation along the villous axis was enabled. Bioprinting technology has also been gaining increasing interest and it has been extensively used for tissue engineering applications. The development of intestinal models is no exception, and authors have been using cell-printing techniques to obtain more realistic models [22, 32, 33]. Kim et al. [32] developed a villus structure using a cellladen collagen bioink cross-linked with tannic acid. In the meantime, authors developed the same villus structure and seeded the intestinal epithelial cells Caco-2 on the scaffold, instead of embedding them. It was observed that cell viability of the cell-laden model was higher than the control and that the expression of MUC17, E-cadherin and alkaline phosphatase was significantly higher and earlier in the cell-laden structure. Nevertheless, beside the better results of the cell-laden model, the control used in the study is more realistic and more physiologically representative, since the intestinal epithelium in vivo sits on the lamina propria. Later, Kim et al. [22] also developed a 3D villi intestinal model using a cell-printing process. The authors used two collagen-based bioinks, one laden with Caco-2 cells for the epithelium layer and one laden with human umbilical vein endothelial cells (HUVECs) for the blood capillary structure [22]. HUVECs were able to form a capillary network inside the hydrogel, and the authors concluded that the 3D villi model containing epithelium and the capillary network demonstrated higher cell growth, expression of enzymes and MUC17 when comparing with the 3D model without capillaries or the 2D model [22]. Here, like in the previous article, but differently from all the others studies already mentioned, epithelial cells were also embedded in the collagen bioink. As stated before, this approach may not be so interesting because, on the human small intestine, epithelial cells are on top of the intestinal lamina propria, forming the epithelium, and not entrapped in the ECM. Nevertheless, the use of HUVECs inside the hydrogel is a very interesting approach, since it mimics the capillaries inside the intestinal lamina propria, which are the last barrier that compounds have to cross to reach systemic circulation and that can have an impact in the absorption process. More recently, Castan˜o et al. [26] proposed a simple and moldless fabrication technique consisting in reaction-diffusion mediated photolithography to fabricate 3D microstructures with complex geometries on poly(ethylene glycol) (PEG) hydrogels. Authors claim that controlling some fabrication parameters such as oxygen diffusion/depletion, distance to light source and the exposure dose, they can define the dimensions and geometry of the microstructures. Caco-2 cells were seeded on these scaffolds and they were able to form a barrier with more physiologic TEER values, and the permeability of FITC-dextran 4 kDa using this model was significantly higher than that obtained using

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the standard Caco-2 model, which is a good indicator, since the permeability of paracellularly absorbed compounds in Caco-2 models is underestimated. It is true that in vivo architecture is very important when mimicking the intestinal villi, giving important cues for cells and altering their behavior in vitro, making them act more physiologically. Studies using this architecture show that cells respond to it and provide very interesting proofs of concept. Nevertheless, not all laboratories and researchers have the facilities and the know-how to produce these complex 3D structures. So, if the aim is to develop a new 3D intestinal model that can be more representative of the in vivo environment, but also simple enough so that anyone can replicate and test their compounds, even if they are not experts in cell culture and biofabrication, other approaches can be useful. Intestinal models that take into account the fact that the model should be a tridimensional structure, representing the different layers of the intestine, but do not represent the complex villi architecture, can be a promising strategy regarding intestinal models aimed to test the permeability of compounds. Some works describe the development of 3D intestinal models without representing the villous structures. Li et al. [20] developed a 3D intestinal model comprised of a collagen layer embedded with fibroblasts, which play an important role in the growth of epithelial cells, and immunocytes with epithelial cells (Caco-2 and HT29-MTX) seeded on top. They were able to get a model with decreased expression of P-glycoprotein (P-gp) and increased expression of breast cancer resistance protein (BCPR) that related more with in vivo levels, and the permeability correlation was also improved [20]. Pereira et al. [17] also developed a 3D intestinal model using Matrigel with intestinal myofibroblasts embedded, and Caco-2 and HT29-MTX cells on top (Fig. 3). The authors showed that the fibroblasts were able to remodel the ECM by production of fibronectin, and the model allowed the efficient prediction of insulin absorption [17]. The downside of this model in terms of reproducibility is regarding the use of Matrigel, since the composition of this product is not completely defined and there is batch-to-batch variation, which can be a problem when trying to reproduce the model, allowing for variability in the results obtained. Dosch et al. [27] tested three different hydrogel scaffolds to develop an intestinal model: alginate, l-pNIPAM (hydrogel composed of 9% N-Isopropylacrylamide (NIPAM), 1% Laponite and 90% water) and l-pNIPAM-co-DMAc (hydrogel composed of 7.83% NIPAM, 1.17% N-N 0 -dimethyl acrylamide (DMAc), 1% Laponite and 90% water). Authors claimed that cells layered on l-pNIPAM hydrogel scaffolds were capable of forming villus-like structures. Nevertheless, these cells formed multilayers, which is not representative of the intestinal epithelium, where cells form a monolayer [27]. Others have also developed intestinal models using different materials and showed interesting results. Maddel and colleagues [33] used Novogel and NovoGen bioprinter, whereas De Gregorio et al. [34] tested a collagen based hydrogel and a cell-synthesized stromal equivalent. The latter claimed that Caco-2 cells were able to differentiate in all four types of intestinal epithelial cells and produced a higher quantity of basal lamina components

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Fig. 3 (A) Schematic representation of the triple-culture 3D model. (B) Hematoxylin and eosin (H&E)stained paraffin section showing myofibroblasts sustaining epithelial Caco-2 and HT29-MTX cells over them. (C) Fibronectin expression in a paraffin section of the triple-culture 3D model. Fibronectin was labeled with Alexa-Fluor 488 (green) and nucleus with DAPI (blue). Reprinted with permission from Pereira C, et al., Dissecting stromal-epithelial interactions in a 3D in vitro cellularized intestinal model for permeability studies. Biomaterials 2015;56:36–45.

like laminin, collagen type IV and hyaluronic acid when using the stromal equivalent, whereas fibroblasts were able to secrete more ECM components [34]. Nevertheless, when Caco-2 cells were cultured on top of this hydrogel, they formed multilayers, which can be a disadvantage as it is not representative of the in vivo situation. Later, the same authors used the stromal equivalent, and patterned it mimicking the intestinal crypt-villus architecture. Using gelatin porous microbeads, they cultivated human intestinal subepithelial myofibroblasts (ISEMFs) and then transferred to a maturation chamber to allow their molding into a disc-shaped construct [35]. A polymethylmethacrylate (PMMA) holed grid was used to obtain the micropatterned surface. The authors observed that the patterned stroma increased the absorbing surface area, epithelial proliferation rate and density of microvilli. In addition, it was able to induce changes in mucus production, polarization and tightness of the epithelial cells. When considering the gastrointestinal tract, it is also important to understand that, in a disease state, the cellular environment can change and, consequently, this can have an impact in different processes, including absorption. Taking this into account, Leonard et al. [36] developed a 3D model to understand the impact of inflammation in the intestinal barrier. The model was composed by immunocytes and dendritic cells embedded in a collagen layer with Caco-2 cells seeded on top. Regarding the TEER values, authors observed that, upon stimulation of the cultures using IL-1ß, there was a decrease in the values to about 80% of the non-stimulated control values, which was similar to what was observed in the 2D Caco-2 monoculture model. It was also observed that the release of IL-8 protein

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into the medium in response to the stimulation was significantly higher in the 3D co-culture in comparison to the Caco-2 2D monoculture. Regarding the 2D monoculture model, it was observed an increased permeability of the paracellularly transported dye fluorescein sodium in the state of inflammation, which agrees with the decrease in the TEER values. Although the authors did not test, it would be interesting to assess the permeability in the 3D model and to use other components, namely the ones with transcellular transport, to better understand the role of inflammation in the absorption process. Beside hydrogels, another common practice is to use decellularized scaffolds to develop the in vitro models [28, 29]. Generally, these scaffolds are obtained using intestinal tissues from animals, normally from porcine origin. The fact that the tissues are obtained from species different from the human can be a disadvantage, because of cross-species differences. Pucsh et al. [28] co-cultured Caco-2 cells with primary human microvascular endothelial cells (hMECs) on decellularized porcine jejunal segments in a dynamic bioreactor. The authors concluded that Caco-2 cells resembled more the normal primary enterocytes when comparing to the static Caco-2 model. Nevertheless, Caco-2 formed multilayers when seeded on the decellularized scaffold. The permeability of substances that have a low permeability coefficient was also enhanced within the dynamic cultures [28]. Li and colleagues [29] also used decellularized porcine small intestine to seed Caco-2 cells, with the aim of obtaining a faster differentiation of these cells, comparing to the 21 days needed when they are seeded on Transwell inserts. The authors verified that, after 7 days, the characteristics of the cells were comparable to the 21-day Caco-2 cells, and that the permeability values of the tested compounds also showed a strong correlation to the cells cultured for 21 days. However, the advantage authors refer to is only in terms of time needed to obtain the model and not the model’s performance itself.

2.2 Organoids When talking about 3D structures comprising the cell types of a specific organ in a selforganized tissue-like structure, we are talking about organoids. Indeed, there are several definitions of organoids given by the experts in the field. For instance, Fatehullah et al., defined organoid as “an in vitro 3D cellular cluster derived exclusively from primary tissue, embryonic stem cells, or induced pluripotent stem cells, capable of self-renewal and selforganization, and exhibiting similar organ functionality as the tissue of origin” [37]. Definitely, all the definitions revolve around the use of several different cell types derived from stem cells that are capable of self-renewal and self-organization, with the ability to have similar behavior to that of the original tissue [38, 39]. As research in organoids increases, a nomenclature was given to distinguish different types of organoids based on the cellular source and composition. In Table 1, we described the three types of organoids and their main characteristics. A tissue-derived organoid can be classified as epithelial, if only grown from epithelial cells, or epithelial mesenchymal, if

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Table 1 Organoids nomenclature based on their composition and cellular source. Organoid type

Characteristics

References

Tissue-derived epithelial

One cell layer of epithelium; generated from differentiated human or animal tissue; extremely expansible; formed in less than 2 weeks; no mesenchymal presence One epithelial layer underlying mesenchyme; generated from tissue; grown in air-liquid interface; mimics malignant histology in vitro; no human model One epithelial layer underlying mesenchyme; generated by differentiation of induced pluripotent or embryonic stem cells; fruitful in vivo transplantation; present fetal characteristics

[40–43]

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[44–46]

[47–50]

grown with mesenchyme and also epithelial cells. The third organoid classification relies on the use of human embryonic or induced pluripotent stem cells (PSCs), more comparable to the fetal than the adult tissue [51], which after transplantation acquires a closer adult-like tissue behavior [52]. In this chapter, the term organoid will be used to refer to tissue-derived structures, with or without mesenchyme, and also PSC-derived structures. In 2009, intestinal organoids appeared and spurred research around these 3D structures. Sato et al. built, from single sorted leucine-rich repeat-containing G-protein coupled receptor 5 stem cells (Lgr5+), a mice intestinal crypt-villus with differentiated cell types and self-organization capability when cultured with laminin-rich Matrigel and growth factors, a model that could be cultured for up to 8 months without losing its features [40]. The authors demonstrated the creation of an intestinal organoid with more than 40 crypts consisting of a single cell layer composed by Paneth and stem cells located at the crypt base, fully polarized enterocytes with mature brush borders, goblet cells, and enteroendocrine cells. In the same year, Ootani et al. used a different methodology through the use of an air-liquid interface exposed to a 3D culture matrix based on collagen and, using primary mouse intestinal fragments, epithelium-mesenchymal organoids were created [44]. This extremely proliferative epithelium, settled by an epithelial polarized cell layer composed by absorptive enterocytes, goblet, enteroendocrine and Paneth cells exhibited structures as microvilli, mucus granules, endocrine granules and intracellular connections of junctional complexes. In the ECM, myofibroblasts were found expressing α-smooth muscle actin and the authors found that these organoids could be cultured for up to 1 year because of the addition of growth factors that simulated Wnt signaling, which represents what happens in vivo in the intestinal crypt [44]. As already mentioned regarding the other intestinal 3D systems, the use of growth factors and ECM during in vitro culture is very important. The organoid development

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and differentiation into 3D-like structures, their own self-renewal and the prolonged time in culture is only possible due to the cell source and the ECM support, as well as the use of growth factors. Matrigel and collagen are the preferred ECM used when culturing organoids, however, other matrix components based on alginate, elastin, laminin, PEG or agarose hydrogels are also used [40, 53–57]. The use of organoids has several advantages, such as the specificity to a person, there are no genetic alterations, similar structures and functions as in vivo are found, and it is possible to culture them for long time-periods [58, 59]. However, the use of embryonic stem cells may cause ethical problems and we are talking about of a time and costly consuming method. Moreover, there is still a lack of vasculature or physical environment, features that leave organoids a step behind to what happens in the living environment. After all, the research has evolved to combine, for example, mechanical forces or fluid flows in order to stimulate cell differentiation or to promote the influx of oxygen and nutrients [60, 61]. Gjorevski and colleagues used Matrigel as control and synthetic PEG hydrogels, varying their composition and mechanical environment to create an intestinal organoid with properties as the ones found in vivo [60]. The authors proposed stiff and dynamic matrices based on PEG (sPEG—mechanically static PEG; dPEG— mechanical dynamic PEG) and, varying the ratios of sPEG and dPEG they were able to control the softening profile of the ECM and cell differentiation and proliferation. The results presented by the authors demonstrated that the 3D microenvironment, ECM constitution and mechanical stimuli have a great influence on the organoid formation, proliferation and self-organization. Thus, these factors may be considered when testing the intestinal permeability of molecules. Since the organoid expansion and behavior is conditioned by the microenvironment surrounding, leading to the mimicry or not of the real situation, the studies performed on those systems will be also conditioned by those factors. Onozato et al. developed an intestinal organoid to study the drug pharmacokinetics, an important feature still little studied in the organoids field. PSCs were induced to differentiate into the mid-intestine to form spheroids, and these were later differentiated into intestinal organoids using small molecules as 5-aza-29-deoxycytidine (5-aza) and N-[(3,5-difluorophenyl) acetyl]-L-alanyl-2-phenyl-1,1-dime-thylethyl ester-glycine (DAPT) [62]. Data revealed an expression of intestinal markers, and pharmacokineticrelated genes as MUC2, chromogranin A, CDX2 and villin, occludin, SLC15A/PEPT1, ABCB1/MDR1 and ABCG2/BCRP were found in the polarized lumen of the epithelium. Since this organoid expressed the functional ABCB1/MDR1 efflux transporter, it could be used in drug development and screening, or even for mechanistic studies to assess the effect of ABCB1/MDR1 in the intestinal permeability [62]. Organoids lead to the creation of an artificial in vitro complex niche, with reflect in vivo features as the highly proliferative and differentiated cell population, and the related physical and chemical environment of the tissue. The recreation of the in vivo situation and, due to

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the fact that the phenotype and karyotype remain unchangeable over time, the development of organoids and here, specially, intestinal organoids, started to be used in the study of intestinal diseases [63–65], cancer [66–68] or in regenerative medicine [69–71]. The epithelial barrier integrity and regulation is a key driver to study several diseases of the intestinal tract, as the case of inflammatory bowel disease and colorectal cancer. Thus, using different stimuli, as the exposure to cytokines, may help to understand the barrier function in disease scenarios and, hopefully, to better understand the disease. Bardenbacher et al. presented a method to analyze the function of the epithelial barrier in terms of permeability and molecular regulation on the 3D intestinal organoids through the use of microscopic techniques and bioinformatics analysis of 3D image data [72]. In this study, the organoids were created as previously described [40, 41], aiming for the translation from 2D to 3D cultures to better understand the interferon gamma (IFN-γ) role in the barrier function. Lucifer yellow was used to study the intestinal permeability using an outside to inside approach, where Lucifer yellow fluorescence was quantified by densitometric analysis [72]. Results demonstrated that the permeability of the barrier increased when induced by IFN-γ, which promoted the cleavage of the claudins of the intestinal barrier (Fig. 4). To confirm the claudins’ destruction and analyze other effects of IFN-γ in other proteins of the epithelial barrier with, 3D reconstruction of confocal z-stack images were used for quantification, based on a Hessian analysis [73]. However, other techniques to study the permeability of molecules, as microinjection or inside to outside approach, are used in the intestinal organoids field. Hill et al. developed a real-time technique for measuring the intestinal permeability in 3D organoids using fluorescently-labeled dextran as model drug (Fig. 5). The procedure was based on the microinjection of the model drug into the apical epithelial surface of the organoid lumen and, at the same time, it was imaged on an inverted microscope, and images were acquired at determined timepoints in order to quantify the fluorescence as mean of permeability [74]. This technique can be employed to deliver pharmacologic agents and with the possibility to measure their permeability in real-time. Leslie et al. used a similar procedure, this time to study the pathogenesis of the Clostridium difficile (CD), a leading cause of infectious nosocomial diarrhea. Human intestinal organoids (HIO) derived from PSCs were used and CD was microinjected in their lumen with the aim of studying CD colonization in the human intestine [75]. Due to the use of organoids with polarized epithelium and paracellular barrier function, it was possible to identify a toxin responsible for the intestinal epithelial barrier disruption in an infectious case. This may be a tool to identify other epithelial barrier effects when studying different types of diseases and testing the permeability of molecules. The study of the intestinal TJ is also related with the intestinal permeability, since TJ modulate the passage of ions and molecules by regulating the paracellular permeability. Organoids constituted by intestinal stem cells, absorptive enterocytes, Paneth cells and secretory goblet cells presented similar morphogenic and morphometric characteristics

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Fig. 4 IFN-γ induces breakdown of the epithelial barrier in mouse intestinal organoids. (A) Intestinal organoids from IFN-γR2WT and IFN-γR2ΔIEC were cultured in the absence (PBS) and presence of IFN-γ for 48 h. Upon addition of Lucifer yellow (457 Da), confocal fluorescent images were captured every 5 min for 70 min. Representative images at time point 0 min and 70 min are shown (green: Lucifer yellow, size bar¼ 50 μm). (B) Fluorescence was determined in the organoid lumen and outside of the organoid for the organoids depicted in (A). Relative intensity values were calculated (fluorescence inside/outside + inside) and are shown for each time point. (C) Upper panel: only organoids with a diameter of 80 20 μm were used. Mean values of the respective organoid diameters are shown (IFN-γR2WT: n ¼ 30, IFN-γR2ΔIEC: n ¼29). The mean diameter values were not significantly different between the different groups by one-way ANOVA. Lower panel: The relative (Continued)

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Fig. 4, Cont'd organoid permeability of IFN-γR2WT and IFN-γR2ΔIEC organoids was determined by dividing the relative fluorescence intensities obtained after 70 min by the minimal relative fluorescence intensities measured during the observation period. Each bar represents mean values, measured in 10 organoids derived from three independent experiments (IFN-γR2WT: n ¼30, IFN-γ R2ΔIEC: n ¼29). Only in IFN-γR2WT organoids intestinal permeability is significantly increased after stimulation with IFN-γ. *** is equal to a p-value < .001 in the Student’s t-test. (D) Organoids were cultured for 48 h in the presence of IFN-γ, IFN-γ + zVAD or left untreated (control). Upon addition of LY, intraluminal LY fluorescence was measured (1 h, 5 min intervals). IFN-γ-induced uptake of LY was not reduced by zVAD. (E) After the experiment in (D), organoids were fixed and apoptosis was visualized by staining for activated effector caspase-8. Only very few apoptotic cells (arrows) are visible under the different conditions (size bar ¼ 10 μm). Reprinted with permission from Bardenbacher M, et al., Permeability analyses and three dimensional imaging of interferon gammainduced barrier disintegration in intestinal organoids. Stem Cell Res 2019;35:101383.

to those found in vivo, and were used by Pearce et al. to study the heterogeneity of the TJ among different cell types [76]. This study allowed to understand that different intestinal cell types display different varieties and levels of TJ’ proteins and that this difference had a significant influence in the macromolecular permeability (Table 2). Watson and colleagues tried a different approach by creating intestinal organoids from PSCs [47, 77] and engrafted them in vivo, in mice [50]. The organoids presented columnar intestinal epithelium surrounded by supporting mesenchyme and took about 35 days to differentiate. Then, they were embedded into collagen type I and engrafted under the mice kidney, and allowed to maturate and grow during 6 weeks. The organoids highly increased their size, presented vasculature, crypt-villus architecture, submucosal layers as lamina propria, muscularis mucosa and submucosa, and smooth muscle layers, and the authors claimed to have found a mature intestinal tissue after the engrafting [50]. The epithelial barrier functionality of the engrafted organoid was tested by a permeability assay through the injection of FITC-dextran. To quantify the permeability, all the blood was collected 30 min and 4 h after injection, and the serum fluorescence was quantified. Data revealed an increase of FITC-dextran permeability overtime, reaching 10% of permeability after 4 h, indicating a good functioning of the epithelial barrier. Intestinal organoids can be a powerful tool in drug screening in a patient’s disease, since it is possible to obtain results in a short time period and, mainly, because it is a patient-directed study. Thus, the identification of the drugs’ response when testing them in the organoids may help to find a new course of treatment, and in vitro tumor organoids can also predict the chemosensitivity of several targeted molecules. Organoids carrying mutations can be useful on the screening of molecules to develop targeted therapies against those mutations. This can lead to an advanced drug screening process and targeted therapy [58, 59]. Organoids can also be used in the biobanks’ development with the aim of analyzing genetic profiling, drug response or possible biomarkers. 20 consecutive colorectal

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Fig. 5 Representative images of a tissue-derived HIO injected with FITC-dextran. (A) Image demonstrating the positioning of the glass microcapillary just prior to insertion into the HIO and microinjection. (B) Brightfield image of a tissue-derived human intestinal organoid after microinjection of FITC-dextran. Note that the fluorescence signal is apparent even without the use of a specific filter set. This coloration aids microinjection precision. Scale bar represents 10 μm. Reprinted with permission from Hill DR, et al., Real-time measurement of epithelial barrier permeability in human intestinal organoids. J Vis Exp 2017;130:e56960.

carcinoma patients were used to collect healthy and tumor tissues for the development organoids for high-throughput drug screening [78] (Fig. 6). Studies as this one can fill the gap between genetics and cancer, and improve the patient treatment through the use of organoids for drug testing and efficacy, obtaining a more personalized therapy.

2.3 Gut-on-a-chip models As mentioned before, Caco-2 in vitro models in Transwells present some disadvantages, such as prolongued culture time, static environment, 2D architecture and the lack of Table 2 Relative permeability of organoids as well as mRNA and protein expression of TJ proteins. Cldn2 Cldn7 Ocln ZO1 Cdh1 Organoid Permeability mRNA Protein mRNA Protein mRNA Protein mRNA Protein mRNA Protein

ISC ENT GOB PAN

+ ++++ ++++ +

++++ ++ + +++

++++ +
++

+ ++++ ++ +++

++ ++++ +++ ++++

++++ +++ +++ ++++

++ +
+++

+++ +++ +++ ++++

+++ ++ ++ +

+++ ++ ++ ++++

++++ +++ ++++ ++++


¼ no expression, + ¼ low, ++ ¼ modest, +++ ¼ high, ++++ ¼ very high expression. Cldn, claudin; ENT, enterocyte; GOB, goblet cell; ISC, intestinal stem cell; PAN, Paneth cell; TJ, tight junction; ZO-1, tight junction protein-1. Reprinted with permission from Pearce SC, et al., Marked differences in tight junction composition and macromolecular permeability among different intestinal cell types. BMC Biol 2018;16(1):19.

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Fig. 6 Derivation of organoids from primary tissue. (A) Overview of the procedure. A total of 22 tumor organoids and 19 normal control organoids were derived and analyzed by exome-sequencing, RNA expression analysis and high-throughput drug screening. To determine the concordance between tumor organoids and primary tumor, DNA from the primary tumor was also isolated. (B) Organoids architecture resembles primary tumor epithelium. H&E staining of primary tumor and the tumor organoids derived of these. A feature of most organoids is the presence of one or more lumens, resembling the tubular structures of the primary tumor (e.g., P8 and P19b). Tumors devoid of lumen give rise to compact organoids without lumen (P19a). Scale bar, 100 μM. Reprinted with permission from van de Wetering M, et al., Prospective derivation of a living organoid biobank of colorectal cancer patients. Cell 2015;161(4):933–945.

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important cell types and mechanisms of the intestinal tissues. So, traditional models fail to mimic complex functions, which are important for drug development and disease research. However, microfluidics has revealed to be a very promising strategy to mimic the human intestine. In this way, human gut-on-a-chip has been developed as a more cost-effective, simple, fast and physiologically relevant model. This model presents 3D architecture, constant flow and mimics the mechanically active microenvironment of the living intestine, which facilitates cell-cell and cell-matrix interactions [79]. In fact, if a system is able to replicate external cellular conditions, the cells will replicate and behave more closely to the in vivo conditions. For instance, for clinical applications, flow is useful to study drug dynamics, delivery of molecules, as well as transport and absorption mechanisms that are not present in static models. Besides, with this type of systems, low volumes are required and so, can, for example, concentrate cell-derived factors. Plus, the user has easy access to cell-conditioned medium and can control the shape and interconnectivity of cell compartments [79]. The chip can be produced by semiconductor industry, using computer microchip manufacturing, as litography techniques. The mold is typically generated from silicon, glass or plastic, and parameters such as mechanical compression or cyclic stress can be regulated and adjusted as desired in order to better mimic the intestinal environment [80]. In general, gut-on-a-chip consists in four chambers (Fig. 7): two central chambers where cells can be seeded and the culture medium circulates, and two vacuum chambers (one on each lateral side) responsible for controlling external factors, such as the membrane strech and the peristaltic motion simulation [81, 82]. Some devices are inclusively used to study the intestinal drug metabolism as well as the interaction of the intestine with nanoparticles, nutrients or host-microbe. For instance, it was observed that a 3D porous membrane in a gut-on-a-chip model potenciates the upregulation of villin and sucrase-isomaltase genes, responsible for intestinal cell differentiation, as well as CYP3A4 and CYP2C9, two cytochrome enzymes related to the intestinal metabolism [83, 84]. Despite forming a planar epithelial monolayer, Caco-2 cells can form undulating structures when in a gut-on-a-chip after 2 days. In addition, it was found that the gut-on-a-chip model promotes the Caco-2 cells’ differentiation in the four main types (absorptive, mucus-secretory and enteroendocrine cells present in the villus regions and, Paneth cells present in the crypt region). These authors also showed that the mechanically active intestine stimulates CYP3A4 activity, glucose reuptake and mucus production, which is known to be absent in static Caco-2 models [85]. Here, once again it was shown the importance of external factors like mechanical forces in getting closer to the final idea, the in vivo human intestine. Also, it was reported that fluidic shear contributes to the higher expression of occludin and ZO-1 proteins in Caco-2 cells, which are responsible for forming TJ and, consequently, an increased in the TEER [84]. Plus, authors considered the mechanical strain and created a biomimetic gut-on-a-chip that mimics the structure, physiology and microflora of the intestine.

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Fig. 7 The mechanically active human Gut Chip. (A) Human villus intestinal epithelium and vascular endothelium are lined on opposite sides of a flexible porous membrane under fluid flow and peristalsis-like strains. A zoom-in schematic shows the intestinal microenvironment undergoing complex crosstalk between commensal gut microbiome, bacterial pathogens, and immune cells in parenchymal and vascular channels, respectively. (B) Villus morphogenesis of human Caco-2 intestinal epithelium in the Gut Chip under physiologically controlled motions and flow. (C) An overlaid image of the co-culture of green fluorescent protein-labeled Escherichia coli and microengineered villi in the Gut Chip. Bars¼ 50mm. Reprinted with permission from Bein A, et al., Microfluidic organ-on-a-chip models of human intestine. Cell Mol Gastroenterol Hepatol 2018;5(4):659–668.

Interestingly, under these conditions, the co-culture of Caco-2 cells with Lactobacillus rhamnosus GG (LGG) demonstrated high TEER values over time. It is also described that external factors accelerate cell differentiation within 3 days [81]. However, this can be a disadvantage, since tighter TJ do not facilitate the transport of molecules and so do not mimic a real situation. On the other hand, others authors verified that cyclic strain enhances the cell differentiation, formation of 3D villi-like structures and paracellular permeability, without changing the TEER values in the monolayer, meaning that mechanical forces act directly in paracellular mechanisms [81]. All in all, the TEER values are not always concordant across literature and this might be due to the current TEER devices being designed to fit into conventional Transwell culture plates and not into microfluidic devices. As curisosity, embedded electrodes inside microfluidic organ chips

3D intestinal models

have already been designed. The creators showed in both human lung airway chip and human gut chip the utility of this embedded electrodes to assess formation and disruption of the barrier function. In this way, the users can have real-time results through a noninvasive approach [86]. Alternatively, some authors choose to incorporate HIO through induced PSCs to an amenable model for study. To form HIO, these authors supplemented induced PSCs with Actividin A, CHIR99021, noggin, fibroblast growth factor 4 and epidermal growth factor to induce endoderm and hindgut formation, as to maintain the organoid formation [87]. Besides having the four intestinal cell types of diferentiation and the characteristic crypt-villus 3D architecture, this model has the advantage of responding to exogenous stimuli [87]. More recently, the same authors aprimorated the intestineon-a-chip to recapitulate the dynamic host-microbiota interface. In this sense, they recreated a system with hipoxia gradient across an endothelial-epithelial interface. The authors claim that the intestinal barrier integrity is actually better under these conditions. This system includes microscale oxygen sensors, Caco-2 cells or human organoid derived epithelial cells, housed in an anaerobic chamber, and enables to keep different commensal aerobic and anaerobic microorganisms in direct contact with patient-derived human intestinal epithelium for 5 days [88]. Donald Ingber and colleagues also fabricated a human small intestine-on-a-chip (intestine chip) composed by 3D organoids of epithelial cells isolated from healthy regions of intestinal biopsies. According to their transcriptomic analysis, this intestine chip is closer to the human duodenum than duodenal organoids [89]. This can be a personalized strategy for each individual, a way to understand specific mechanisms and therapies, and to get more reliable results. These organoids are dissociated and cultured on a porous membrane. This system is also characterized by having human intestinal microvascular endothelium cultured in a parallel microchannel under flow and cyclic deformation. Plus, as advantage, this intestine-on-a-chip is useful for metabolism, nutrition, infection and drug pharmacokinetics studies, considering the easiness to quantify nutrient digestion, mucus secretion and establish a intestinal barrier function [89]. Meanwhile, there is a commercial version available. Mimetas, the organ-on-a-chip company, created the Organoplate that combines microfluidics to a standard well plate. This allows the user the freedom to decide which ECM, cell types, co-culture and perfusion to use, according to each interest. In addition, this company recently developed a methodology to assess the barrier integrity using 40 leak-tight, polarized epithelial gut tubes, exposed to an ECM. The method was recreated in a microfluidic environment, expressing polarization and transporters. This can be a way to simulate a pathological situation like dysfunction of epithelial barriers, or drug-induced toxicity, which can reduce cellular life conditions and drug development. The epithelial barrier disruption condition is characterized for having an increased paracellular permeability. The authors claim to be a method of friendly use, sensible, and with real-time analysis (Fig. 8) [90].

407

(B)

(A)

400 µm x

z

(D)

ECM gel

220 µm

y Phaseguides

(E)

Cells

(K)

Caco-tube with flow

Flow ECM gel channel

(I) Day 0

(C)

y

ECM gel

(M)

(G)

(N)

(H)

(O)

(I)

Day 7

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(F)

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(J)

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Fig. 8 Overview of the method for modeling intestinal tubules in the OrganoPlate platform. (A) Photograph of the bottom of an OrganoPlate showing 40 microfluidic channel networks with inlay showing the top view of the 384-well plate format device; (B) Zoom-in on a single microfluidic channel network comprising three channels that join in the center. (C, E, G, I) Horizontal projection and (D, F, H, J) vertical cross section of center region for subsequent steps in establishing the gut model. (C, D) An ECM gel (light gray) is patterned by two phaseguides (dark gray), (E, F) culture medium is introduced in the two lanes adjacent to the ECM gel, one of which comprises cells. (G, H) Cells are allowed to settle against the ECM gel surface by placing the plate on its side. (I, J) Upon application of flow, cells form a confluent layer lining the channel and gel surfaces, resulting in a tubular shape. (K) 3D artist impression of the center of a chip comprising a tubule, an ECM gel and a perfusion lane; two phaseguides (white bars) are present that define the three distinct lanes in the central channel. The tubule has a lumen at its apical side that is perfused. (L–P) Phase-contrast images of the formation of the tubular structure at day 0, 1, 4, 7, and 11, respectively. Scale bars are 100 μm. Reprinted with permission from Trietsch SJ, et al., Membrane-free culture and real-time barrier integrity assessment of perfused intestinal epithelium tubes. Nat Commun 2017;8(1):262.

3D intestinal models

The following table sums up some examples of 3D intestinal models developed using microfluidics (Table 3). Basically, it is possible to combine different approaches and aprimorate the intestinal models with, for example, immune cells, intestinal microbiota, organoids or mechanical forces, to approximate in the maximum possible way to an in vivo situation [91]. Although, further studies are developing systems encouraging the integration of the liver, for instance, since it is important for reproducing the first-pass metabolism. In this sense, different organs-on-a-chip have been developed, in order to connect them and mimic the whole system present in the body and to have more trustable results [92]. In addition, with a more complex system, it is possible to have a broader perspective of the drugs’ effects, meaning that it is possible to check potential side effects in different organs [93, 94]. The future goal is to include in the whole system and cells from patients like induced PSCs. In this way, treatments can be uniquely designed for each patient [95, 96]. It was already reported a functional coupling between five microphysiological systems (MPS) to validate the transport, metabolism, blood-brain barrier permeability, and toxicity of drugs. These MPS represent the organs responsible for absorption, metabolism and clearance of the drugs, as the jejunum, liver and kidney, respectively, as well as skeletal muscle and neurovascular models. After testing terfenadine, trimethylamine (TMA)

Table 3 Examples of in vitro 3D intestinal models for permeability screening using microfluidics. Model

Compounds

Advantages/applications

Reference

Caco-2

Caco-2 and LGG

PDMS and ECM

HIO

PDMS and matrigel

Gut tubules Caco-2, B. fragilis

ECM PDMS, oxygen sensor, collagen, matrigel Porous membrane

Promotes cellular differentiation and intestinal metabolism Caco-2 differentiation in four cell types Mimics the shape and density of the human intestinal villi Mimic the influence of the microflora in the barrier functions Able to respond to exogenous stimuli To study the barrier integrity It has both anaerobic and aerobic environment Personalized strategy for each individual

[83]

Caco-2

Porous nitrocellulose membrane and collagen type I Porous PDMS membrane Collagen scaffold

Caco-2

Intestinal organoids

[85] [84] [81]

[87] [90] [88] [89]

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and vitamin D3, the authors obtained results showing an agreement between organ-ona-chip and clinical data. Plus, it was discovered that trimethylamine-N-oxide (TMAO) was able to cross the blood-brain-barrier. Here, it is obvious the potential of these MPS to use sequential fluid transport to study different organs’ effects on downstream organs and to evaluate drugs’ toxicity and absorption, distribution, metabolism and excretion (ADME) (Fig. 9) [97].





(A)

Goblet Cell Paneth Cell

Intestinal Epithelial Cell Enteroendocrine Cell

(B)

Hepatocytes

Kupffer Cell

Endothelial Cell

Stellate Cell

Collagen

Media

(C)

Proximal Tubule Endothelial Cell

Pericyte

Neuron

HUVEC

Endothelial Cell

Astrocyte

(D)

Fig. 9 Schematic representations of the four organ systems used for functional coupling. (A) The intestinal module is constructed in Transwells from primary jejunum enteroids. Test agents are applied in the apical compartment h1i. The media collected in the basolateral compartment h2i is used to add to the liver. (B) Media from the jejunum intestine basolateral compartment h2i is perfused as a 1:3 jejunum/naïve liver media into the influx port of the SQL SAL liver model h3i. Efflux media is collected h4i and used to add to two downstream organ models. (C) The vascularized kidney proximal tubule module is a two lumen, dual perfusion system. For the vascular compartment, jejunum/liver-conditioned media h4i is diluted 1:2 or 1:4 with naïve EGM-2 media and then perfused into the influx port h5i to collect effluent from the proximal tubule at h6i. In parallel with perfusion through the vascular compartment, the proximal tubule compartment is perfused with naïve DMEM/F12 PTEC media h6i for effluent collection. (D) The blood-brain barrier with neurovascular unit (NVU) is constructed in a membrane-separated, twochambered microfluidic device. The brain-derived endothelial vascular compartment is perfused at the influx port h7i with jejunum/liver-conditioned media h4i diluted 1:4 with naïve EGM-2 media. The effluent is collected at the efflux port h7i. In parallel with perfusion through the vascular compartment, the neuronal cell compartment is perfused with naïve EBM-2 media at the influx port h8i for effluent collection at h8i. Reprinted with permission from Vernetti L, et al., Functional coupling of human microphysiology systems: intestine, liver, kidney proximal tubule, blood-brain barrier and skeletal muscle. Sci Rep 2017;7:42296.

3D intestinal models

3. Limitations and future perspectives Until this point, a review of enhanced intestinal cellular models with focus on intestinal permeability was performed. It was possible to observe that 3D models that arose in recent years bring innumerous advantages comparing to the 2D models that are normally used. These advantages are not only regarding drug development, but these models also allow a better understanding of cells’ behavior in vivo. Nevertheless, there are still limitations regarding these models. Starting with the 3D multilayered models, it was possible to understand that, in recent years, a lot of development has been made. Different research groups are trying to build better models and that can more precisely mimic the in vivo environment, and the main aim is to get as close as possible to what is observed in vivo. It is obvious that models that are able to replicate the intestinal architecture, mimicking the villus structures, the crypts or even both, are more realistic and comparable to the human small intestine, leading to the differentiation of epithelial cells along the crypt-villus axis, which correlates to what is observed in vivo. It is also true that this cellular differentiation can have an impact in the behavior of cells regarding absorption processes [18, 23, 24]. The presence of villi itself is also important when evaluating the absorption of compounds, since it enhances the absorption area [26]. Nevertheless, the fact that different groups are trying different approaches and using different techniques to mimic the small intestine architecture can be a drawback regarding permeability testing. The fact that there is not a standard 3D model, as it is the case of the Caco-2 model, is probably hampering their use in permeability assays. In fact, in most articles regarding development of new formulations, authors still use the Caco-2 model. This also happens because some of the 3D models that have been discussed here are not easy to obtain, requiring specific equipment and know-how. Truth is, scientists that are focused on drug development normally prefer platforms that are easy to obtain and handle. As so, the 3D models without the architecture may be easier to adopt by scientists focused on drug development when it comes to test the permeability of their compounds. Nevertheless, these models are still not perfect and there is room for improvement. Considering the amount of studies of the last years regarding this topic, it is expectable that new and more complex models will arise in the future, always with the aim of trying to better replicate the in vivo situation and to have results that are more accurate and reliable. The advantage of using 3D organoids systems is tremendous. Being constituted by several cell types, presenting ECM and acquiring the complexity and structure from a determined organ, are the main reasons why organoids are so similar to what is found in in vivo organs. Additionally, these systems can be generated from induced PSCs, which can lead to a patient-specific organ and, consequently, to a patient-specific treatment based on the study of the disease mecanisms in these 3D systems. Drugs can be screened in order to assess its eficacy and effects and, thus, this approach leads to a more specific and eficient treatment, since not everyone respondes in the same way to the same drug. Also, the creation of

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organoids from normal or diseased tissue can help us to better understand and characterize the events that occur during all stages of disease progression and development. Besides the advantages referred above, there are several limitations that keep organoids away from what happens in a living organ. First, the use of induced PSCs is limited and entails ethical problems. Moreover, the use of induced PSCs can generate different cell types of a specific organ, but reaching the in vivo maturity is difficult and a lack of neural, endothelial, or immune cells and physical microenvironment (as fluid flow or peristaltic movements) in the intestinal organoids is still an aspect to consider when comparing to the in vivo situation. Also, it is a time consuming and costly technique that has variations from laboratory to laboratory, from tissue source to tissue source, leading to heterogeneity and reproductive failure. When talking about testing the permeability of molecules in organoids, a more complex process is found. In most of the cases, the molecule has to be injected in the organoid, being a time-consuming and variable technique. Also, the TEER can not be monitored to assess the barrier integrity, however, epithelial proteins and/or imaging techniques can be used to study the barrier function. The same happens with gut-on-a-chip systems, with different groups developing models according to their needs. The main drawback of these systems is the fabrication method, which needs expertise in microengineering process, besides the variability from batch to batch. Also, PDMS is highly used and its high permeability to small and hydrophobic molecules can interfere with the results of the experiment. Some use bioprinters and face a reduced optical transparency, which is a desirable feature for imaging cellular processes. Regarding permeability, the main challenge is maybe the balance between features, to guarantee the proper TEER values. Regarding the future perspectives of gut-on-a-chip systems, some authors are considering to incorporate induced pluripotent stem cell-derived organoids in gut-on-a-chip models, which enhance the opportunities for personalized therapeutics. Others are thinking in integrating important components such as intestinal fibroblasts, immune cells and the enteric nervous system. The value offer of these models is the possibility to mimic the in vivo human intestine by integrating different cell types and external factors (shear stress, mechanical forces, flux) as well as the ability to analyze the complexity of the systems at different growth stages. Either 3D multilayered models, intestinal organoids or gut-on-a-chip systems can offer a good correlation with the in vivo situation compared to the 2D standard systems. The next steps in this field to get closer to the human intestine may go through the combination of knowledge and new approaches to achieve complex models as substitutes to the in vivo experiments.

4. Conclusion In the recent years, great advances have been achieved in the development of 3D systems that better mimic the in vivo environment. Although still far from the human organ

3D intestinal models

complexity and functioning, it is possible to find 3D intestinal models encompassing the ECM, the physical environment or even the structure, and several cell types of an organ. Thus, the research performed in these 3D systems, as the screening of the intestinal permeability of molecules or compounds, leads to more reliable results, with improved correlation to the in vivo situation, which does not happen with 2D systems, since they are very limited. Therefore, assessing drug absorption in these systems may have a positive and significant impact on the development of preclinical drugs or formulations, decreasing the time of research and the experimental animals used, since more reliable data can be obtained.

Acknowledgments This chapter was financed by the project NORTE-01-0145-FEDER-000012 by Norte Portugal Regional Operational Programme (NORTE 2020), and COMPETE 2020—Operacional Programme for Competitiveness and Internationalisation (POCI), under the PORTUGAL 2020 Partnership Agreement, through the FEDER—Fundo Europeu de Desenvolvimento Regional, and by Portuguese funds through FCT— Fundac¸a˜o para a Ci^encia e a Tecnologia/Ministerio da Ci^encia, Tecnologia e Ensino Superior in the framework of the project "Institute for Research and Innovation in Health Sciences" UID/BIM/04293/2019. Andreia Almeida, Cla´udia Azevedo and Maria Helena Macedo would like to thank to Fundac¸a˜o para a Ci^encia e a Tecnologia (FCT), Portugal for financial support (SFRH/BD/118721/2016, SFRH/BD/ 117598/2016 and SFRH/BD/131587/2017, respectively).

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CHAPTER 14

In vitro relevant information for the assessment of nanoparticles for oral drug administration
Alonsoc,d María Victoria Lozanoa,b, Manuel J. Santander-Ortegaa,b, María Jose a

Cellular Neurobiology and Molecular Chemistry of the Central Nervous System Group, Faculty of Pharmacy, University of Castilla-La Mancha (UCLM), Albacete, Spain b Regional Centre of Biomedical Research (CRIB), University of Castilla-La Mancha (UCLM), Albacete, Spain c Center for Research in Molecular and Chronic Diseases (CIMUS), Campus Vida, University of Santiago de Compostela, Santiago de Compostela, Spain d Department of Pharmaceutics and Pharmaceutical Technology, School of Pharmacy, Campus Vida, University of Santiago de Compostela, Santiago de Compostela, Spain

1. General considerations of oral drug delivery The oral route is the most attractive modality of administration of drugs, distinguished by its high patient acceptance and convenience. Drug absorption by the oral route is favored for most of the type I and type II drugs of the Biopharmaceutical Classification System (BCS). Nevertheless, this is not the case for some other drugs with poor stability or low permeability, whose low bioavailability is the main reason for their failure in clinical trials [1]. Insulin is one of the most representative examples, however, an increasing number of marketed biological drugs also exhibit these characteristics, such as those indicated for bowel diseases (IBD) and other autoimmune pathologies [2, 3]. The identification of these limitations at an early stage of the costly drug development process will allow tackling these hurdles through the design of adequate delivery strategies. Nanotechnology has led to an increasing number of nanocarrier formulations, which may have the potential to overcome the biopharmaceutical limitations of those molecules [4]. Nanocarriers that are specifically designed for oral drug delivery are expected to have the capacity to confront harsh environmental conditions, such as salinity, acidity and presence of enzymes, which may be deleterious in terms of physical and chemical stability. In addition, they have to come across the mucus layer that protects the intestinal epithelium; otherwise, they will be cleared away from the intestinal lumen [5]. The intestinal epithelium itself is a complex structure of different cells, with enzymatic activity, efflux pumps and cell trafficking, which also represents a great challenge for the transport of nanocarriers and/or their associated drugs. Therefore, the success of the development of new oral medicines involves the identification of the most potent formulations [6]. Because this process is long and costly, large efforts are focused on the development Nanotechnology for Oral Drug Delivery https://doi.org/10.1016/B978-0-12-818038-9.00014-4

© 2020 Elsevier Inc. All rights reserved.

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of methods that help to identify those with the best biopharmaceutical properties and efficiency [7]. In vitro models for the evaluation of oral formulations have a prominent position on the development process as they are cost-effective, allow high throughput assays and provide a certain resemblance to the in vivo conditions [8, 9]. The main challenge of in vitro models is fulfilling biomimicry, by matching the complexity of the biological environment with their setup. Ideally, these models would help to characterize nanoformulations and predict their performance in the human body [10]. Obviously, there is not a unique model that gathers all of the characteristics of the oral barrier. Instead, there are different in vitro models of increasing degree of complexity that allow the understanding of the mechanistic behavior of nanocarriers in biological media. They dissect the effect of parameters like pH or enzymes among others in a controlled fashion, so the adequate selection of the in vitro models and the combination of the provided information is of great value for researchers. Altogether, the role of in vitro testing methods on the oral medicines development process is undeniable; hence, their optimization and standardization is expected to have a great impact on the costs of the overall preclinical testing. This chapter provides insights into the in vitro techniques for the characterization of nanoparticles intended to oral delivery. Initially, it details the features of the oral route, focusing on the barriers that it represents for drug delivery. As these barriers are the ones that the in vitro models would need to replicate for its reliability, a careful explanation of the composition of biological fluids of the intestinal tract and the mucosal barrier will be given. It is within this context that the material-biological interaction becomes important, as they determine the biological performance of the nanosystems. Then, an overview of the different nanotechnological approaches at the material-biological interface will be presented. Understanding the performance of the interface in relation to biological media can help us to identify the gaps and challenges that will promote the future approaches in nanoparticulate drug delivery. The in vitro characterization techniques will include those studying (i) the stability in gastrointestinal fluids, (ii) the mucoadhesion/mucodiffusion, and (iii) the interaction and transport across the intestinal epithelium. A recent area of research are the microphysiological systems containing gut-on-a-chip, as a combination of the lab-on-a-chip technology with tissue culture methods [11, 12]. All these methodologies are aimed toward reducing costs, increasing reproducibility and provide promising prototypes for the subsequent in vivo studies [13].

2. Barriers to oral drug delivery The human gastrointestinal track (GIT) is constituted by specialized tissues and accessory organs for the digestion and absorption of nutrients and water. It exhibits large surface area, covered by a protective layer of mucus, and harsh conditions like acid pH values,

In vitro relevant information

presence of enzymes, bile salts and microbes. In addition, the intestinal epithelium is a very selective barrier and has a role on the regulation of the physiologic homeostasis. Therefore, the oral barrier could be considered as a physical barrier, constituted by the mucus layer and intestinal epithelium, but it is also a functional barrier, due to the effect of GI enzymes, efflux pumps and components that mediate the transport and biotransformation of molecules.

2.1 Gastrointestinal juices and microbiota Once in the stomach, nanoparticles will be exposed to low pH (1–3) and a significant medium ionic strength (I
100 mM). Gastric fluids in fasted and fed states contain surface-active molecules that will have a great affinity for the nanoparticle/water interface. These molecules include enzymes, polysaccharides or phospholipids, which can promote both the aggregation of the formulation and its enzymatic degradation [14]. In the duodenum, the first section of the small intestine, the formulation will be exposed to a gradual increase in the medium pH (up to 6.5), to the bile salts secreted from the liver via the gall bladder and to a cocktail of enzymes and co-enzymes produced by the pancreas. In addition, the transition from the stomach to the intestine entails an increase in the medium salinity (it can rise up to 140 mM) [15]. Another component of the gut that has gained increasing attention is the human microbiota. There is compelling evidence suggesting the implications of gut microbiota on human health and autoimmune or inflammatory diseases such as IBD, multiple sclerosis and cancer [16]. In this sense, the alteration of gut microbiota may have an impact on the intestinal permeability and mucus layer thickness [17]. Bacterial metabolism is an issue to consider in drug delivery, as it can determine the stability of drugs in the intestinal lumen due to the large number of bacteria in the intestine [18]. A recent work by Ohno et al. has reported a curcumin nanoparticle formulation that suppressed the development of dextran sulfate sodium-induced colitis, which might be related to the regulation of the microbial structure [19]. These evidences suggest that the role of microbiota on the development of nanocarriers should be carefully evaluated [20].

2.2 Mucus Mucus is a protective layer that lubricates the intestinal epithelium and controls the access of foreign pathogens. Briefly, this mucus layer is a viscoelastic gel with a complex composition of water (
95%) and mucins (2–5%) as major components, and minor components like cells and cellular debris, proteins, DNA, salts and lipids, among others [21, 22]. Mucins are megamolecules (Mw 0.5–40 MDa) secreted by the goblet cells and the submucosal glands (Brunner´s glands in the case of the intestinal tissue). These megamolecules present a molecular backbone with a dense coverage of glycans and cysteine-rich domains [23–25]. This heterogeneous composition allows the formation of reversible linkages between the

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mucin fibers through both low affinity non-covalent and stronger disulfide bonds [21, 26]. This negatively charged mesh presents pore sizes ranging from 50 nm up to 2 μm [23]. Additionally, the mucus structure depends on the diet, age and pathological state [25].

2.3 Cellular barriers The absorption of nutrients occurs, mainly, at the small intestine, which is a tubular structure constituted by four different layers: mucosa, submucosa, muscularis externa and serosa [27]. The mucosa increases its surface area by folding into villi and microvilli to maximize absorption. Villi and microvilli size decreases from the proximal portion of the duodenum to the distal ileum [28]. The oral mucosa is an organized structure specialized on the absorption of nutrients. Thus, most of the cells of the intestinal epithelium are columnar epithelial cells called enterocytes (nearly 90%) [29]. The complexity and efficiency of the intestinal epithelium requires other type of cells, such as mucus-secreting goblet cells, paneth cells that secrete antimicrobial lysozymes, as well as enteroendocrine cells and tuft cells for the regulation of digestion and absorption [30]. Peyer’s patches are 1% of the cells that constitute the brush border, and they belong to the lymphatic system. Lymphoid follicles are covered by the follicle-associated epithelium (FAE), which contains microfold cells (M cells). These cells belong to the intestinal innate immune system and translocate particles, antigens, bacteria and viruses by transcytosis [31]. The mucosa of the intestine is exposed to harsh conditions and, hence, the epithelium renewal is paramount for its well function throughout life. This is supported by small intestine stem cells residing inside the crypts, which are deep tubular glands located between villi. The renewal process of the small intestinal epithelium occurs, approximately, every 4–5 days, yielding a high renovation rate. Epithelial cells are immersed in a dynamic environment of intestinal peristaltic waves, blood flow and chemical gradients. Other components of the cellular barrier are the commensal bacteria that inhabit the intestinal lumen, ranging from 103 to 109 organisms per mL of luminal content from the duodenum to the ileum [32].

3. Understanding the biological-material interface In general, the ultimate goal of nanoparticulate oral drug delivery systems is to reach the intestine. Nanoparticles unique properties come from their miniaturization and massive surface/mass ratio [33, 34]. This favors their interaction with the components of the complex biological environment. As DLVO stated in the late 1940s, the interaction between two approaching nanoparticles (VT) can be expressed as an equilibrium between the repulsive potential created by the ionisable surface groups of the nanoparticle (VE) and the attractive potential created by the hydrophobic Van der Waals forces (VA) [35, 36]. Modulation of energetic barriers involved in the colloidal stability of the

In vitro relevant information

formulation by the presence of electrolytes and macromolecules in the administration medium will be decisive in the behavior of the nanosystems in biological media. The complexity of the gastrointestinal environment is a great challenge for nanoparticle efficiency, as they have to face chemical and enzymatic degradation, physicochemical instability processes, entrapment in the intestinal mucus and demanding intracellular trafficking. Nevertheless, such complexity can also be considered as a source of opportunities to promote the interaction with the biological environment and to potentiate the performance of the systems. This can be achieved by modulating the composition of nanoparticles and the formulation method to adapt their physicochemical properties to the intestinal environment. The careful design of the nanoparticle composition according to the intestinal environment will lead to formulations with more favorable biological behavior [37].

3.1 Electrolytes and pH Variations of the electrolytes concentration along the gastrointestinal tract will result in the modification of the electrical double layer (EDL) thickness of the nanoparticles and the subsequent change in the total potential of interaction. The increase of the medium salinity will compact the EDL, and the VE will suffer a sharper decay that can result in the cancellation of the energetic barrier that keeps the formulation stable. Under this scenario, all the collisions between particles will result in aggregation. The minimum electrolyte concentration required to reach an energetic barrier equal to zero is known as critical coagulation concentration (CCC). Determination of the specific CCC value of a formulation is a key aspect to predict its behavior after the oral intake. Those formulations with a moderate CCC value (50–100 mM) will reach the stomach intact but will collapse there. In the intestinal tract, the nanoparticles will face two different scenarios: those with CCC values in the 100–150 mM range will aggregate in the intestines, and those with CCC values higher than 200 mM will be stable. The CCC of each formulation will depend on its specific surface physicochemical properties. Fuchs introduced an experimental factor, known as Fuchs or stability factor (W) that considers the collision efficacy between two approaching nanoparticles. This is the probability of collapse after the collision of the nanoparticles that form the formulation [38]. Afterwards, Verwey and Overbeek, two of the authors of the DLVO theory, connected the experimental W factor with the total potential of interaction between two approaching nanoparticles as follows: Z ∞ ðVT =kT Þ e W ¼ 2a 2 dH 0 ð2a + H Þ Where a is the radio of the particles, k the Boltzmann’s constant, T the temperature and H the distance between the nanoparticles. This experimental approach allowed the

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calculation of the specific electrolyte concentration that can lead to the premature aggregation of a formulation. Most of the colloidal systems used in drug delivery have ionisable surface groups. Under this scenario, the medium pH and electrolytes have a similar effect on the stability of the formulation. At pH values close to the isoelectric point of the nanoparticles, the reduction of the charged surface groups minimizes the repulsive potential of interaction VE, leading to the aggregation of the formulation. Analogously, at pH values far from the nanoparticle isoelectric point, the charged surface groups can stabilize the formulation. The colloidal behavior of lyophobic systems can be fairly well predicted by the DLVO theory, but, as stated by Overbeek himself, it fails for lyophilic formulations [39]. In this sense, DLVO cannot explain the colloidal behavior of formulations that contain chitosan, or other well-known hydrophilic polymers widely used for oral drug delivery, such as poly(ethylene glycol) (PEG) [40–42]. To understand the colloidal behavior of these formulations, it is necessary to consider the effect of the hydration forces on the interaction of the nanoparticles with the surrounding medium, a non-DLVO structural force that can stabilize hydrophilic formulations in aqueous media. The repulsive potential of interaction created by this structural force can be expresses as follows [43]: Vh ðH Þ ¼ πaðNA Ch ce Þλ2 eðH=λÞ Where NA is the Avogadro number, Ch the hydration constant and ce the electrolyte concentration. As can be extracted from the expression, the intensity of this repulsive force depends on both the hydrophilic character of the nanoparticle surface and the electrolyte concentration of the media. Hydration forces explain the unexpected behavior observed for chitosan nanocapsules under intestinal conditions. Indeed, chitosan nanocapsules, which have been proposed as oral drug delivery vehicles, were found to have an acceptable stability in biological media [22]. The role of hydration forces on the colloidal stability of chitosan nanocapsules was confirmed by the analysis of their colloidal stability in intestinal buffers prepared at different ionic strengths [41, 44]. While the formulations were colloidally unstable at low ionic strengths, the increase of the electrolyte concentration of the medium resulted in stabilization of the formulations, even when the nanocapsules had null surface charge. This effect can be attributed to the hydrophilic character of chitosan. The combination of a hydrophilic surface and the presence of highly hydrated ions results in the formation of a shell of highly structured molecules of water. This structural force depends on the hydrophilic character of the formulation (Ch) and the concentration of the hydrated ions, and presents a relevant effect on the stability of hydrophilic formulations. Decoration of the surface of nanoparticles with neutral poly-oxyethylene (PEO) derivatives has been the most widely used strategy to increase the hydrophilicity of colloidal drug delivery systems [33, 34, 42, 45]. Studies from our group have shown that

In vitro relevant information

PEG-coated nanoformulations may have different responses to the electrolyte concentration depending on the coating density. At low coating densities, the presence of PEG results in a reduction of the ζ-potential of the formulation, which cannot be compensated by the slight increase of the hydrophilic character of the formulation, and leads to instability of the system (reduction of CCC) [46]. By increasing the coating density, the system evolves from pure electrostatic controlled stability to an electrosteric one and the CCC starts to increase [33, 42]. Under this scenario, the formulation becomes more hydrophilic and sensitive to hydration forces, but its stability still depends on the electrolyte concentration. Finally, by increasing even more the coating density, the formulation should become stable at any salt concentration. The absence of aggregation phenomena by electrolytes is indicative of a steric stabilization of the formulation [33, 34]. This nonDLVO stabilization process presents two main contributions. Thus, two new terms should be added to VT, i.e., an osmotic contribution (when the surface layer of PEO derivatives overlap) and a volume restriction contribution (due to the compression of the PEO derivative layers when the particles are getting closer than the polymer layer thickness). These forces (Vosm and Vvr, respectively) can be enthalpically or entropically driven, and are mainly dependent on the nature of the solvent-PEO derivative interactions [34]. This means that the specific contribution of Vosm and Vvr to the total potential of interaction will be inherent to each specific PEO derivative.

3.2 Enzymes and active surface molecules The gastrointestinal track is highly specialized in the digestion to facilitate the absorption of nutrients across the intestinal epithelium. This task is performed by a great number of enzymes and active surface molecules present in the gastrointestinal juices. These macromolecules have great affinity for the nanoparticle surface, so it is necessary to modulate their interaction with the formulation to avoid its premature aggregation or degradation, and maximize the success of the therapy. In line with the previous section, a further question is whether the approaches used to minimize the effect of the pH and electrolytes of the GIT on the colloidal stability of the formulations can be exploited to modulate the interaction of the system with the macromolecules of these media. This hypothesis would minimize the number of excipients, increasing the robustness of the formulation. The total energy of an adsorption process can be expressed as ΔGAds ¼ Δ HAds  TΔSAds. Based on the endothermic nature of the adsorption, it could be assumed that the driving force of macromolecule adsorption must be entropic and comes from the release of (i) water molecules (hydrophobic dehydration), (ii) counterions, or (iii) changes on the macromolecule structure (conformational entropy). The adsorption of macromolecules onto colloidal systems is mainly driven by the dehydration of the hydrophobic

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surfaces [47]. Therefore, the use of hydrophilic surface agents can reduce their interaction with proteins and other macromolecules. Hydroxyethyl-derivatives of poly(amino acids) such as poly-asparagine and polyglutamine, have shown promising results as surface agents to reduce the interaction of the colloidal systems with the surrounding macromolecules [48, 49]. Similar results were obtained by Peng et al. by using albumin as coating agent to reduce the non-specific interactions of the formulation with biomacromolecules [50]. However, the proteolytic environment of the gastrointestinal tract could limit the applicability of these approaches to the oral route. Other hydrophilic alternatives are polyoxazolines, hydroxypropyl derivatives of methacrylamides or polysaccharides such as chitosan, hyaluronic acid, dextran or heparin [40, 41, 44, 51]. The use of zwitterionic polymers with functional groups like phosphorylcholine, sulfobetaine or carboxybetaine can also reduce the interaction of the formulation with the surrounding macromolecules [51–53]. Interestingly, for zwitterionic coatings, the increase in the ionic strength up to physiological conditions was reported to result in the improvement of the stealth properties [54]. These mechanisms are also present in the stealth properties obtained with PEG derivatives [55]. The incorporation of these polymers on the colloids neutralizes the surface charge of the formulation, and subsequently the potential release of counterions during the adsorption process [53]. Additionally, the hydrophilic nature of these derivatives will also minimize the adsorption of the macromolecules by the reduction of the hydrophobic dehydration contribution to ΔGads. In parallel with these energetic barriers, the presence of a polymeric shell will avoid the deposit of macromolecules onto the nanoparticle surface through steric and osmotic repulsions [55, 56]. These stealth properties will depend on the coating density and the conformation of the PEG chains on the surface of the nanoparticles, as well as on their molecular weight (Fig. 1) [55, 57, 58]. The stealth properties of the PEG coating are usually reached under a brushed conformation at high coating densities [45, 55, 59]. However, for some colloidal systems, the macromolecules-repulsion

Fig. 1 Different conformations of the PEG chains as a function of the coating density of the nanoparticle and the energetic barriers created as a function of the specific conformation.

In vitro relevant information

effect of PEG is also observed at low coating densities [60, 61]. Under these conditions, the capacity of macromolecules to penetrate the PEG layer leads to an osmotic repulsion that can end up in the change of the PEG chain from a pancake, mushroom or loop to a brush conformation (process known as osmotic constraints). PEG molecular weight will affect the arrangement of the polymer onto the nanoparticle surface, where the use of high molecular weight polymers can result in the increase of the intra- and inter-molecular repulsions and the formation of more heterogeneous and less dense coatings [40, 44, 58]. A more detailed discussion about the effect of the PEG coatings on the interaction of nanoparticles with biological macromolecules can be found elsewhere [55–57].

3.3 Mucus Over the last decades, significant efforts have been directed to the design of mucoadhesive nanoparticles with the idea of improving the delivery and transport of drugs across the mucosa. However, the results of these efforts led to the conclusion that a strong interaction with the mucus may actually interfere with the enhanced penetration. This is due to the turnover of the outer loose layer of mucus that covers the intestinal tract. As a result, researchers proposed the necessity of achieving a mucoadhesion/mucodiffusion balance that facilitates both the interaction and the penetration of particles across mucus [62]. Several types of raw materials, either natural or synthetic, have been used to promote the adhesion of the formulations to the intestinal mucus. Chitosan, poly-acrylic acid, alginates, cellulose derivatives, such as hydroxyethyl-, hydroxypropyl-, and carboxymethylcellulose (HEC, HPC, and CMC respectively) or hydroxypropyl methylcellulose (HPMC) are some examples of mucoadhesive materials [63]. The mechanisms behind the retention of the nanoparticles in the mucus include hydrogen bounds, hydrophobic or electrostatic interactions, as well as direct entanglement with mucins [22, 24, 64]. An attempt to increase mucoadhesive properties of the materials has been their chemical modification to include thiol groups. This improves the retention of the formulation in the mucosal tissue thanks to the formation of disulfide bonds between the modified polymer and the cysteine-rich domains of mucins [65, 66]. On the other hand, in order to facilitate the permeation of the particles across the mucus, some authors proposed the use of mucus disrupting nanoparticles. This can be achieved either by the immobilization of mucolytic agents in the nanoparticles, or by the chemical modification of mucoadhesive polymers, such as poly-acrylic acid with mucolytic polymers or enzymes, like papain or bromelain. This has shown a clear effect on the mucus penetration capacity of the tested formulations [67, 68]. An alternative approach to facilitate the mucodiffusion of the particles led to the conception of mucus penetrating particles (MPP) by Hanes et al. [26, 69, 70]. In this case,

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unlike the mucus disrupting approach, the passage of the nanoparticles across the mucus layer is achieved through the mucus pores without disrupting its barrier properties [25, 26, 69, 71–73]. The diffusion of the particles across the water pore maze requires minimizing any interaction with the mucin fibers. This is feasible by obtaining a dense coating of nanoparticles with PEG or PEG derivatives with a Mw 10 KDa (>10 chains/ 100 nm2) [45, 74, 75]. This is in line with the above-mentioned alternative to avoid the interaction of the formulation with the biomolecules present in the gastrointestinal juices. However, due to the complex nature of the mucin fibers, higher coating densities are required to obtain nanoparticles able to efficiently diffuse across the mucus blanket [74, 75]. Other alternatives to PEG are the use of zwitterionic or ζ-potential-shift polymers [76]. Irrespective of the agreement on the potential of MPP, it has been argued that adhesion might also be of interest for the treatment of local pathologies of the intestinal mucosa [63].

3.4 Intestinal epithelium As explained in the previous section, overcoming the oral barrier firstly involves the interaction with the mucus layer but, afterwards, with the intestinal epithelium as well. Disclosing the mechanisms of cell trafficking for nanoparticles is crucial for the design of efficient nanocarriers. Broadly speaking, the intestinal epithelium is a barrier constituted mainly by enterocytes whose cellular membrane confers integrity to the epithelium. This epithelium is supported by the presence of tight junctions that attach epithelial cells. Therefore, the mechanism of transport between cells, known as paracellular transport, is normally impaired for nanoparticles. In addition, tight junctions represent only a small fraction of the whole surface of the epithelia. Thus, the interaction of the nanoparticles with the cells is mostly mediated by the transcellular route [77]. Endocytosis is an active transport mechanism that can be divided into phagocytosis and pinocytosis [78]. Phagocytosis is a process of M cells and phagocytic immune cells, such as macrophages or dendritic cells that mediates the uptake of large particles. Besides, pinocytosis involves the uptake of fluids and solutes. Pinocytosis occurs in all types of cells and can be divided into clathrin-mediated endocytosis and clathrin-independent processes [79]. The interaction of nanoparticles with the intestinal epithelium is a complex process that depends on different factors, such as the nanoparticle size, geometry, surface charge, surface composition and functionalization. In this sense, the work recently published by Banerjee et al. highlighted the role that nanoparticle shape has on the intestinal uptake process. The authors used the triple co-culture model of Caco-2/HT-29/Raji-B lymphocytes and showed that the rod-shape nanoparticles were more efficiently taken up by the cells, compared to the spheres and discs [80]. With regard to the effect of the size, those same authors have shown a size-dependent cellular transport with the

In vitro relevant information

highest rate for those particles with the smallest sizes (50 and 200 nm). The surface charge of the nanocarriers is obviously related to their surface composition [81], and it can affect their interaction with the epithelium, their stability and mucodiffusion, as it will be explained below. This section will focus on the results of in vitro studies analyzing solely the nanoparticle interaction with the epithelium. Since the origin of chitosan nanoparticles and nanocapsules [82–84], this biomaterial has been extensively used for the preparation of transmucosal drug delivery vehicles [85]. The interest in this biomaterial has relied on its bioadhesive and penetration enhancing properties. However, in vitro studies have suggested that the strong bioadhesive properties of chitosan may shield its interaction with the epithelium [86]. Therefore, as indicated above, a balance between mucoadhesion/mucodiffusion might be necessary to achieve the necessary performance. In this sense, N-(2-hydroxypropyl) methacrylamide polymer was used to cover N-trimethylated chitosan cores associating insulin. The rationale behind this approach was to confer the system with a hydrophilic coating for mucodiffusion that, when properly detached, would lead to N-trimethylated chitosan cores to efficiently interact with the cell membrane. This core-shell nanostructure achieved higher insulin transport across a mucus-secreting cell culture model compared to the uncoated system, and transiently opened tight junctions in a reversible way, facilitating the transcellular passage of insulin [87]. Targeting to goblet cells can be accomplished by modifying the nanoparticles with the peptide CSKSSDYQC (CSK). This was observed when CSK modified trimethyl chitosan coated nanoparticles were tested on the mucusenterocyte model. The conjugation to the CSK peptide led to higher uptake of the systems due to the faster opening of the TJ, and to the involvement of clathrin-mediated endocytosis and other mechanisms like caveolae-mediated endocytosis and macropynocitosis of the unmodified system [88]. Another approach was followed by Verma et al. when they developed multi-layered chitosan nanoparticles decorated with vitamin B12 to promote the interaction with the intrinsic factor receptor located at the enterocytes. This surface modification increased insulin delivery while maintaining the tight junction opening, as observed by the transepithelial electric resistance (TEER) measurements of the enterocyte model [89]. The cell trafficking of nanoparticles can lead to degradation of the cargo molecule by the effect of the lysosomal enzymes. A recent work by Fan et al. addresses this issue by taking advantage of the bile acid pathway, proposing deoxycholic acid-modified chitosan nanoparticles as a novel strategy for insulin delivery [90]. These examples illustrate how the functional modifications can be used to maximize the interaction of surface tailored nanoparticles with the intestinal epithelium. PLGA is a polymer widely used in drug delivery for its biocompatibility, biodegradability and for being considered as safe for human use by the Food and Drug Administration (FDA) and European Medicines Agency (EMA) [91]. There is a strong evidence of the utility of PLGA for oral administration of proteins [92], cyclosporin [93] and anticancer drugs [94]. In addition, early studies have shown that PEGylation would

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be critical for preventing their stability and mucodiffusion [26, 69, 70]. In this respect, Song et al. compared PEG lengths from 2000 to 5000 Da and observed minor variations in terms of their cell uptake [95]. The surface properties of PLGA nanoparticles can be modified by the inclusion of polymers such as N-trimethyl chitosan chloride (TMC). The coating with TMC significantly increased the permeability coefficient of insulin compared to the uncoated system, which was suggested to be due to the opening of the tight junctions, as observed by the decreased TEER values [96]. PLGA nanoparticles have also been functionalized with different targeting ligands in order to increase their interaction with the intestinal epithelium. For example, targeting the Na+-coupled organic cation/carnitine transporter 2 OCTN2 (SLC22A5) can be achieved by conjugating carnitine to the surface of the nanoparticles [97]. The results indicated that there was a maximum carnitine density for optimal uptake, a fact that has been observed for other functionalised nanocarriers where high ligand density provokes low particle uptake [98]. Similarly, PLGA nanoparticles have been functionalized with lectins, which specifically interact with glycosylated membrane components [99]. The modification of PLGA nanoparticles with wheat germ agglutinin (WGA) has led to a significant increase of the nanoparticle uptake mediated by EGF receptor-mediated endocytosis [100]. Another approach to enhance the interaction of PLA-PEG nanoparticles with the intestinal epithelium has relied in their functionalization with the neonatal Fc receptor (FcRn). Fc-targeted PLA-PEG nanoparticles loaded with exenatide were significantly transported across the intestinal barrier compared to the unmodified particles, exploiting the natural transport mechanisms of IgG in Caco-2 cells [101]. This strategy enhanced the transport of exenatide loaded in Fc-modified PEG-PLGA nanoparticles, showing a faster and more pronounced uptake when the systems were functionalized [102]. Finally, engineering nanoparticles for M-cell uptake can be achieved by chemically modifying PLGA nanoparticles with mannosamine. Modified nanoparticles strongly interacted with the cells and followed an active mechanism of uptake by receptormediated endocytosis [103]. Polyaminoacids and polypeptides are materials of great interest for drug delivery due to their biocompatibility, chemical versatility and cell-penetrating properties [104]. Considering this, Lollo et al. have recently reported the use of polyarginine nanocapsules as oral carriers of the antitumoral peptide elisidepsin. In this case, polyarginine promoted the interaction with the Caco-2 monolayer, producing a transient decrease of the TEER values that was concentration dependent [105]. A further investigation with polyarginine nanocapsules showed a capacity to increase insulin transport across the Caco-2 monolayer [106]. The arginine-rich sequences of protamine enabled the membrane-translocation properties of the protein, and this has inspired the formulation of protamine nanocapsules for the oral delivery of peptides [107]. Transient TEER values reduction, as observed for polyarginine nanocapsules, and cell uptake mechanisms with higher contribution of caveolae-mediated process compared to clathrin-mediated

In vitro relevant information

endocytosis were observed when the formulations were tested in the enterocyte-like model, and in the inverted FAE model [108]. Phospholipids and lipids are frequent components of nanocarriers for oral drug delivery. They do not only offer increased solubility of poor-soluble drugs, but also modulate the interaction with the intestinal environment, as described in previous sections, and promote cellular drug uptake [109]. The components used and their disposition lead to different lipid based nanocarriers like self-emulsifying drug delivery systems (SEDDSs) or nanostructured lipid carriers. The apparent permeability coefficient of insulin when incorporated in SEDDSs constituted by medium-chain triglycerides was significantly higher than that obtained for the system formed by long-chain triglycerides, and this was observed in both the enterocyte-like and the mucus enterocyte-like models [110]. Nanostructured lipid carriers composed of solid and liquid lipids [111] were evaluated for the transport of saquinavir in the enterocyte-like model and in the FAE model. The results showed that the surfactant content, as well as the size of the carriers, determined the transcytosis mechanism, with caveolae- and clathrin-mediated processes and P-glycoprotein (P-gp) efflux avoidance, thus supporting that the careful design of the nanocarriers is crucial toward the development of efficient delivery systems [112].

4. Nanoparticles for oral drug delivery: In vitro characterization techniques In vitro techniques are powerful tools for the characterization of nanocarriers. They allow the detailed analysis of the different factors and conditions of the intestinal environment that have an effect on the nanoparticles. The studies in simulated gastric/intestinal fluids and mucodiffusion are critical for understanding nanoparticle colloidal stability and their behavior in complex media. On the other hand, the use of cell culture intestinal models enables the screening of the different formulations and the understanding of their mechanistic behavior [113]. Table 1 gathers the advantages and limitations of the different in vitro techniques that are included in this section.

4.1 Stability in gastrointestinal fluids Photon correlation spectroscopy (PCS), also named as dynamic light scattering (DLS), is the most frequently used technique to characterize the stability of colloidal drug delivery systems. DLS instruments analyze the intensity of the light scattered by the nanoparticles via the correlation function. Briefly, a typical correlogram presents an initial plateau (correlation coefficient close to 1) with a length that depends on the particle size. This initial section is followed by a decay in the correlation coefficient down to zero, where the slope of the decay is used to calculate the polydispersity index (PDI) of the formulation. In parallel to the size determination, DLS instruments usually allow the determination of the electrophoretic mobility of the formulation (or ζ-Potential) through the

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Table 1 Main advantages and limitations of in vitro models for nanoparticle characterization.

Colloidal Stability

Mucoadhesion

Method

Advantages

Limitations

DLS

✓ Well-known technology ✓ Fast bench assay ✓ Low amount of sample

NTA

✓ Well-known technology ✓ Fast bench assay ✓ Low amount of sample ✓ Avoids multiple scattering of complex media ✓ Fast bench assay ✓ Low amount of sample

✓ Not physiological conditions ✓ Multiple scattering of complex media ✓ Complications for polydisperse populations ✓ Not physiological conditions ✓ Difficult set-ups for complex media, such as mucus

Mucosa pipe

Mucin film

✓ Fast bench assay ✓ Low amount of sample

DLS

✓ Well-known technology ✓ Fast bench assay ✓ Low amount of sample ✓ Low amount of sample ✓ Allows intensity quantification of the mucus-sample interaction

Mucus rheology

AFM

✓ Low amount of sample ✓ Detailed analysis of the mucus-sample interaction

✓ Not physiological conditions ✓ Recommended fluorescent inner control ✓ Not physiological conditions ✓ Semiquantitative technique ✓ Fluorescent inner control required ✓ Not physiological conditions ✓ Multiple scattering at physiological mucin concentrations ✓ Not physiological conditions ✓ Determines the macroscopic effect of the formulation in the mucus, but not the effect of the mucus in the individual behavior of the particles ✓ Not physiological conditions ✓ Use of non-complex mucin samples

In vitro relevant information

Table 1 Main advantages and limitations of in vitro models for nanoparticle characterization—cont’d

Mucodiffusion

Method

Advantages

Limitations

SANS-NMR

✓ Low amount of sample ✓ Detailed analysis of the effect of nanoparticles in mucin fibers ✓ Fast assay ✓ Quantitative determination of the transport of the nanoparticles across the mucus

✓ Not physiological conditions ✓ Use of non-complex mucin samples

Transwelllike assays

Mucus/ Mucin-filled tube

✓ Quantitative determination of penetration ✓ Low amount of sample

FRAP

✓ Low amount of sample ✓ Quantitative determination of the sample mean diffusion coefficient ✓ Low amount of sample ✓ Quantitative determination of the sample ensemble diffusion coefficient and individual nanoparticle diffusion coefficient ✓ Well stablished model ✓ Morphological and functional characteristics of enterocytes ✓ Good correlation to human absorption for many drugs ✓ Higher paracellular permeability than the enterocyte-like model ✓ Presence of mucus

Tracking

2D cell culture

Enterocytelike model

Mucus enterocyte model

✓ Difficult to control the thickness/uniformity of the mucus layer ✓ Need of steady state in the donor ✓ Issues in pore size threshold selection ✓ Not physiological conditions ✓ Skills to avoid forces that may affect the diffusion profile ✓ Multiple steps protocol ✓ No information about the interaction of individual particles with mucus ✓ Issues with raw mucusnanoparticle analysis ✓ Skills for the proper analysis of the particles trajectories in mucus

✓ Caco-2 heterogeneity, so clones are required ✓ Low paracellular transport ✓ Absence of mucus ✓ Heterogeneity of the mucus layer width ✓ Absence of double mucus layer Continued

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Table 1 Main advantages and limitations of in vitro models for nanoparticle characterization—cont’d

3D cell culture

Method

Advantages

Limitations

Follicleassociated epithelium model

✓ Presence of human M cells ✓ Translocation of nanoparticles and lower TEER values ✓ Use of biomimetic materials ✓ Presence of villi and crypts ✓ Metabolic activity and paracellular transport closer to human intestine ✓ Continuous media renewal ✓ Replicates biophysical forces ✓ Inclusion of vascular component and bacteria

✓ Requires skills in cell culture ✓ Higher proportion of M-cells than in vivo

Engineered intestinal tissues

Microfluidicbased approaches

✓ Need of scaffold for 3D architecture ✓ Scaffold might hinder the transport of molecules

✓ Complicated set-up ✓ Requirement of instruments to elicit flow

laser Doppler velocimetry, phase analysis light scattering (PALS) or a combination of both. A key aspect of this technique is that the medium must allow the free passage of the scattered light from the nanoparticles to the instrument sensor. The presence of a high concentration of macromolecules in the medium results in multiple scattering, where the photons scattered from the nanoparticles are themselves scattered by the macromolecules present in the media, resulting in aberrant correlograms. Nanoparticle tracking analysis (NTA) has recently emerged in the drug delivery field as an alternative to overcome the DLS limitations for size determination in complex media [114, 115]. This technique uses a laser beam to track the Brownian motion of the particles. The scattered light is recorded in a CCD or CMOS camera sensor, and the diffusion coefficient of the formulation is calculated from the mean square displacement of the nanoparticles (more details about this approach can be found in Section 4.2). Additionally, while DLS is based on light intensity analysis, NTA evaluates both light intensity and particle number, which ensures better resolution of polydisperse samples, such as those collected from complex biological-like media [114, 115]. DLS and NTA are also widely used in the analysis of the stability of drug nanocarriers. DLS is a robust technique to monitor the effect of the electrolytes and pH changes on the

In vitro relevant information

colloidal stability of the formulations. However, the studies carried out in the presence of supplemented media with gastrointestinal enzymes and other macromolecules (pepsin and pancreatin are the most used supplements to simulate gastric and intestinal fluids respectively, USP XXIX) should be carefully designed, using the adequate controls for the accurate interpretation of the multiple scattering patterns. Meanwhile, NTA can give relevant information about the particle size populations in the presence of biological molecules. In any case, both techniques are essential to understand the behavior of the nanoparticles in the GIT and guide the fine reformulation of the prototypes. For the interpretation of the DLS/NTA data, it should also be taken into account that gastrointestinal enzymes and other biomolecules can bind to the nanoparticle surface and form either heterogeneous or homogeneous coatings. While the heterogeneous coatings may end up in the massive aggregation of the formulation, a homogenous enzymatic corona can electrostatically stabilize the formulation [116]. This process may not be detected by DLS and NTA as the hydrodynamic mean size may remain stable. In the case of the lipid nanocarriers, the adsorption of pancreatic lipase onto lipid nanocarriers may lead to surface erosion [117, 118]. Such an alteration to the nanocarriers might not lead to a significant change in the particle size or zeta potential and, hence, it is not detected by the DLS or NTA instruments [119, 120]. In these cases, it is mandatory to carry out a complementary in vitro assay to determine the potential digestion of the triglyceride matrix into fatty acids and monoglycerides. A lipolysis assay may simply involve the titration with NaOH of the free fatty acids released to the medium [121–123]. Medium composition is another key aspect of the evaluation of the in vitro colloidal stability. The recommended composition of simulated gastrointestinal fluids has evolved over the years in order to properly mimic the in vivo conditions. Pharmacopeia presents the simplest simulated media, where the simulated gastric fluid (SGF) has a pH of 1.2 and contains pepsin, and the simulated intestinal fluid (SIF) presents a pH of 6.8 and pancreatin (USP XXIX). However, leaders in the field have recommended specific media that simulate the gastric and intestinal fluids in fasted (FaSSGF/FaSSIF) or fed states (FeSSGF/ FeSSIF) [124, 125]. These modifications include the presence of electrolytes, lecithin and other surfactants, or different bile salts to simulate more properly the continuously changing gastrointestinal milieu [126].

4.2 Mucoadhesion and mucodiffusion In vitro techniques to characterize the interaction of colloidal nanoparticles with the mucus can be classified depending on (i) the inclusion of mucus (collected from humans or animals), synthetic mucus or purified mucins and (ii) the disposition of the mucus (immobilized onto a support or in solution). As a function of these variables, in vitro

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techniques go from the study of the nanoparticle-mucin interactions under controlled conditions to the analysis of the effect of the mucus in the Brownian motion of the nanoparticles. Additionally, the development of high-speed cameras has allowed the measurement of the mucodiffusion of individual particles. Bearing this in mind, this section explains the different in vitro techniques used to determine the effect of the formulation characteristics on its interaction with the mucus. They will be summarized as: (i) Handmade Techniques, which make use of simple and low-cost experimental setups; and (ii) Specialized-Equipment Techniques, which require the use of atomic force microscopy (AFM), nuclear magnetic resonance (NMR), strong-anion-exchange (SAX), or high speed cameras. Handmade Techniques are adequate for a fast and early screening of the formulations, while the SpecializedEquipment Techniques usually provide a deeper knowledge of the interaction of the particles with the mucus.

4.2.1 Handmade techniques Analysis of muco-interaction Bernkop-Schn€ urch’s research group has designed different handmade techniques to analyze the interaction of several formulations with different mucosal tissues [127–131]. For example, the group studied the effect of the surface concentration of dSH free groups on the interaction of thiolated-HEC nanoparticles with porcine intestinal mucosa by placing the freshly excised mucosa in a half pipe (with an inclination of 45 degree). Fluorescently-labeled nanoparticles with different surface dSH content were placed in the upper part of the tissue, and subsequently rinsed with saline (1 mL/min; up to 3 h). Interestingly, the fraction of the formulation retained in mucus was proportional to the surface concentration of free dSH groups [65]. A similar assay was performed with thiolated chitosan-decorated liposomes and the presence of the thiomer coating resulted in a clear increase of the liposome retention time in the mucus [130]. The quantification of the immobilization of a formulation onto a glass slide covered by a mucin film has been an alternative and simple way to determine the affinity of the nanoparticles for the mucus [119, 132]. This technique allows a semi-quantitative analysis of the interaction of the fluorescently labeled nanoparticles with mucins. Santalices et al. have also used this approach as an early fast screening of a series of nanocapsules with different polymer shells [119]. DLS has also been used to study the size and ζ-potential evolution of nanocarriers exposed to increasing concentrations of mucin [119]. At a certain mucin concentration, muco-interactive nanocarriers will aggregate and have ζ-potential values closer to the ones of mucins. This technique is limited to the use of mucin concentrations of around 1% w/v, which are far from those of the mucus (2.5–5% w/v). This is a prerequisite for the diffusion of the scattered light to the sensor device.

In vitro relevant information

Mucus

NPs

Analysis of mucodiffusion Mucodiffusion techniques involve the use of diffusion cells, such as Franz cells, Ussing chambers and Transwell assays [133, 134]. The main idea behind these setups relies on the estimation of the flow of nanostructures from a donor-like to an acceptor-like compartment separated by a mucus (or mucin) layer. In the horizontal chambers, mucus sits on a membrane filter, where in the vertical versions, mucus is confined between two membrane filters [135, 136]. The key point of this approach is the pore size of the membranes [133]. Indeed, small pore size is required to retain mucus (or mucin solution) between the donor and the acceptor compartments; however, it may hinder the free passage of the nanoparticles across the filter [133]. Other aspects to consider are the difficult control of the thickness and uniformity of the mucus layer, as well as the necessity of a steady state in the donor-like compartment. Another fast-screening technique involves the analysis of the capacity of fluorescently-labeled nanocarriers to diffuse along a mucus- or mucin-filled tube [21, 119, 129, 137]. As shown in Fig. 2, this analysis may be conducted in two different ways. The first approach is based on the analysis of the penetration profile of the fluorescent nanoparticles in a mucus-filled capillary [119, 138]. In this case, the analysis of the penetration profile can be expressed by fitting of the fluorescence decay versus the distance as IF ¼ A + I0e x/k; where IF is the normalized fluorescence, A is the basal normalized fluorescence, I0 the fluorescence at X¼0 (mucus/formulation interface), and k is the decay constant. This constant can be expressed as k ¼ X1/2/ ln 2, being X1/2 the penetration capacity, that is the thickness (μm) of mucus that allows the transport of 50% of the formulation [119]. Another alternative data analysis is to fit the fluorescence decay to the Fick’s second law as ∂ C/∂ t ¼ D(∂2C/∂2x); where D is the effective diffusion coefficient of the formulation in mucus [138]. The device used in the second approach is a mucus-filled silicon tube. Once filled with the mucus and the formulation, the tube is

Fig. 2 Schema of the different variants of the mucus-filled tube technique to quantify the mucodiffusion capacity of nanoparticles. (A) Use of the fluorescence decay signal from the nanoparticle solution toward the mucus contained in a capillary tube. (A1) The fluorescence decay profile can be used to estimate X1/2. (A2) Second Fick´s law can be used to calculate the diffusion coefficient from the fluorescence decay curve. (B) Diffusion in a mucus-filled silicon-like tube and subsequent analysis of fluorescence in mucus slices.

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incubated at 37 °C under rotation for 24 h to facilitate the penetration of the formulation into the mucus gel. Afterwards, the tube is frozen and cut into slices; the quantification of the fluorescence of each slice allows the determination of the penetration capacity of the corresponding prototype [129]. 4.2.2 Specialized-equipment techniques Specialized-equipment techniques implicate the use of high-performance equipment that provide highly accurate information about the interaction of the formulation with the mucus. As described above, mucus is a heterogeneous gel with a complex composition, where all its components, including trace elements, can dramatically affect its rheology [25]. Bearing this in mind, different research groups have analyzed the effect of the presence of mucoadhesive nanoparticles on the elastic and viscous modulus of the mucus. Mucoadhesive polymers like poly-acrylic acid or chitosan can increase the resistance against deformation of the mucus [139]. This process, known as rheological synergism, correlates with the intensity of the nanoparticle-mucus interaction, and has been used as parameter to determine the mucoadhesive character of different polymers and formulations [26, 140]. AFM has also been adapted to understand the physicochemical mechanisms underlying the interaction of nanocarriers with the mucus [141–143]. Lijima et al. revealed that a shell of poly-methacrylic acid grafted with ethylene glycol (poly(MMA-g-EG)) enhanced the long-range adhesive forces of the formulation with the mucus under intestinal pH due to the swelling of the polymer. Interestingly, under gastric pH, the collapse of the coating shell reduced the polymer-mucin crosslinking, and then the attractive formulation-mucus interaction [143]. A similar AFM experimental setup was used by Cleary et al. to study the interaction of a Pluronic-poly(acrylic) acid nanogel with the intestinal mucosa [141]. These authors described that the interaction of the formulation with the mucin fibers was pH-dependent. AFM studies showed that the electrostatic repulsions between the polymer and mucin hindered the interaction of the nanocarriers with the mucus at pH 5 [141]. Small-angle neutron scattering technique (SANS) or NMR have been used to determine both the structural changes of the mucus mesh in the presence of mucoadhesive polymers and the effect of oral drug delivery formulations on the mobility of the mucin fibers [68, 127]. This technique revealed that the use of poly(acrylic acid) nanoparticles or the same nanoparticles decorated with mucolytic enzymes (papain or bromelain) did not affect the mucus structure at a scale length of 1–400 nm. However, NMR showed that the presence of the mucolytic enzymes clearly increased the diffusion capacity of mucin fibers, most probably due to the disruption of the entanglement stiffness of mucin [68]. Fluorescence recovery after photobleaching (FRAP) has played a key role on the understanding of the barrier properties of different mucosal tissues [144, 145]. FRAP uses fluorescent samples and it is based on the analysis of the fluorescence recovery kinetics of a

In vitro relevant information

delimited area after laser bleaching. The fitting of this kinetics to a mathematical model (most of them based on the Fick´s second law) leads to the ensemble average diffusion coefficient. In this sense, the fluorescence recovery kinetics of mucus containing fluorescent nanoparticles after photobleaching can be used to calculate the diffusion coefficient of the formulation in the mucus [146]. This technique was used to determine the microrheology of human cervical mucus (a well-known mucus model of the intestinal mucosa), and the calculation of the diffusion of proteins, antibodies, viruses and polystyrene nanoparticles. These studies determined that mean pore size of the mucus was around 100 nm [145]. In fact, FRAP has been widely used to obtain the diffusion coefficient of different polymers and nanoparticles in both mucus and mucin solutions [146–149]. This technique has highly contributed to the study of the mucus barrier in drug delivery. Nevertheless, it is important to highlight that FRAP gives information about the ensemble average diffusion coefficient of the formulation, but not of individual particles. The recent advances in CCD and sCMOS sensor technology, software and hardware of fast speed cameras have marked an inflection point in the study of the Brownian motion of nanoparticles. The coupling of this technology to a fluorescence microscope allows the tracking of individual nanoparticles in complex biological media and tissues [150–153]. The experimental limitation of this technique lies in the determination of the nanoparticle displacement from frame to frame of the video. The faster the camera hardware and software (lower lag time), the more accurate the analysis of the Brownian motion of the nanoparticles is. Actually, fast research cameras can record a frame every 1 ms, generating movies of 1000 frames/s. However, the recording speed is also affected by the fluorescence labeling of the formulation, where a minimum excitation time is required to capture the fluorescence coming from nanoparticles. The calculation of the displacement of the nanoparticles from frame to frame can be expressed as:  2   2  Δr ðτÞ ¼ Δx + Δy2 ¼ MSD where Δx and Δy represent the displacement in the x and y directions and MSD the mean squared displacement of the nanoparticle [154]. From the MSD, it is possible to calculate the effective diffusion coefficient (Deff) as: MSD ¼ Deff 4τ where τ is the time scale. In a homogenous medium, the MSD increases with time, as the nanoparticles have more time to move, and the Deff is constant. In a heterogeneous medium, such as the mucus, the situation is slightly different. In this environment, the presence of obstacles or the binding of the particle to mucins results in a subdiffusive transport, and the MSD depends on the time as follows: MSD ¼ Deff 4τα

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where α (the anomalous diffusion exponent) gives information about the nature of the diffusion mode of the nanoparticle in the corresponding biological matrix [155–157]: • Free diffusive nanoparticles (α
1). • Subdiffusive nanoparticles (α