Daniela Patrícia Peneda Pacheco

Development of an Injectable PHBV microparticles-GG Hydrogel

Hybrid System for Tissue Engineering Applications

Dissertação do 2º Ciclo de Estudos Conducente ao Grau de Mestre

em Tecnologia Farmacêutica

Trabalho efetuado sob a orientação de:

Professora Doutora Maria Helena dos Anjos Rodrigues Amaral

Doutora Alexandra Margarida Pinto Marques

Doutor Vítor Manuel Correlo da Silva

Novembro de 2013

É AUTORIZADA A REPRODUÇÃO INTEGRAL DESTA MONOGRAFIA

APENAS PARA EFEITOS DE INVESTIGAÇÃO, MEDIANTE DECLARAÇÃO

ESCRITA DO INTERESSADO, QUE A TAL SE COMPROMETE

i

“Scientific work must not be considered from the point of view of the direct

usefulness of it. It must be done for itself, for the beauty of science”

Marie Curie (1867-1934)

ii

iii

Acknowledgments

I would like to express my gratitude to PhD Professor Helena Amaral, my supervisor from

Faculty of Pharmacy, for the support that she was giving me during this year both

scientifically and academically.

I am completely grateful to my 3B’s Research Group supervisors. To Dr. Alexandra

Marques, for all the brainstorms that she gave me with her knowledge, for always making

me think forward, supporting me in all the ideas and aims. To the other member of the

triad, Dr. Vitor Correlo, I want to say that it was a pleasure to work with him, and I want to

express my indebted to him for sharing his “out of the box” ideas with me, that allowed the

development of a unique work. I am especially thankful for both of them, for the good

relationship, encouragement, availability, persistence, and fruitful discussions.

To Professor Rui L. Reis, for allowing me to conclude this stage of my academic life at his

Research Group, allowing me the access to all the materials and equipment that were

necessary to complete this important step of my academic years. Moreover, he provided

me the contact with great scientists in this field, from whom I learned which is not yet

written in books.

In this 3B’s journey, I had the opportunity to interact with unique people to who I want to

express all my appreciation for the laugh, tears, culture and support. Dr. Praveen Sher I

am deeply grateful for making me overcome the drug delivery systems boundaries with

his tips which definitely made all the difference. To Marta Ondrésik (Márti), I want to thank

her for all the car conversations. Definitely I have never met anyone that justifies the day

to day things with biological pathways from tissues to molecules, I am grateful to you. To

Dr. Manuel Alatorre for always spreading his happiness and smiles through the corridors,

for the early morning rides and also for his knowledge on Physics, Biology and Chemistry

that helped me a lot. To Dr. Marianti Mantha, the Greek that translated English to

Portuguese for making me speak English. And finally, a special thanks to Sebastião van

Uden, who definitely made this year easier and happier, who supported me in the bad and

good moments.

To all the people of 3B’s Research group including technicians and friends, for their good

work environment and friendship.

Um obrigada muito especial à minha família, nomeadamente aos meus pais, que

definitivamente tanto se esforçaram para me dar a mim e aos meus irmãos os alicerces

iv

necessários para chegar até aqui, alcatroando as estradas para que o caminho a

percorrer fosse mais fácil.

v

Resumo

O design e síntese de sistemas eficazes de libertação de fármacos (SLD) é extremamente

importante para a área da engenharia de tecidos e medicina regenerativa (TERM). Estruturas

optimizadas com uma distribuição e uma cinética de liberação bem definidas são pré-requisitos

para uma aplicação clínica segura e eficaz, evitando efeitos colaterais indesejáveis e toxicidade.

Visando o desenvolvimento de SLD de aplicação minimamente invasiva, aliada a uma libertação

sustentada, prolongada e localizada de sinais bioquímicos relevantes para a regeneração do

tecido, foi desenvolvido neste trabalho um SLD injetável bifásico, o qual resulta da combinação de

micropartículas (MPs) de poli(hidroxibutirato-co-hidroxivalerato) (PHBV) (MPs) com um hidrogel de

goma gelana (GG).

A primeira etapa do trabalho apresentado nesta dissertação consistiu na produção de MPs de

PHBV, com incorporação de moléculas modelo de proteínas e glucocorticóides. Para este efeito,

foram incorporadas dentro de MPs de PHBV Albumina de Soro Bovino (BSA) e Dexametasona

(Dex) usando um método de dupla emulsão A/O/A com evaporação do solvente, após algumas

modificações. Visando o controlo adequado do perfil de liberação de moléculas bioactivas, como

fase orgânica foram usadas diferentes proporções de clorofórmio e etanol. Algumas características

do sistema de MPs foram afectadas pela modificação do método, nomeadamente o tamanho das

partículas, bem como a topografia da superfície das mesmas. A influência da presença de etanol

na fase orgânica também foi avaliada em termos de eficácia de incorporação das moléculas

bioactivas, sendo que a sua presença conduziu a uma maior eficácia na incorporação de Dex,

enquanto a incorporação de BSA não foi afectada. Por sua vez, os estudos de libertação in vitro

revelaram que MPs produzidas na ausência de etanol na fase orgânica apresentam um perfil de

libertação mais sustentado quando comparado com as MPs produzidas na presença de etanol.

Estes resultados eram esperados uma vez que a adição de etanol à fase orgânica induziu a

formação de poros maiores na superfície e no núcleo das MPs, proporcionando deste modo a

penetração de água, e consequentemente a libertação mais rápida das moléculas bioactivas.

Após a obtenção das MPs de PHBV com incorporação de moléculas de interesse para a área de

TERM, o objectivo foi definir um sistema que actuaria como um sistema de transporte das MPs,

contribuindo para aumentar o tempo de residência das MPs no local de injecção. Para este efeito,

as diferentes MPs de PHBV foram incorporados nos hidrogéis injectáveis de GG, amplamente

estudado no nosso grupo de investigação para diferentes aplicações de engenharia de tecidos. As

MPs de PHBV incorporadas no hidrogel apresentam-se uniformemente distribuídos na matriz,

demonstrando, também, uma forte integração com o polímero do hidrogel. Como esperado, a

incorporação das MPs nos hidrogéis de GG influenciou fortemente o perfil de libertação, obtendo-

se sistemas de libertação de ordem zero. Embora as propriedades do hidrogel possam ser

modificadas de um modo controlado, com o objectivo de ajustar o perfil de libertação de acordo

com requisitos específicos dos tecidos e da patologia/lesão, o sistema proposto pode, certamente,

ser explorado como uma SLD de longo prazo.

vi

Neste sentido, um SLD bifásico, que permite a libertação de sinais bioquímicos com diferentes

características físico-químicas, bem como a sua libertação localizada, foi desenvolvido com

sucesso e constitui uma ferramenta versátil de modulação de microambientes celulares instrutivos

para aplicações em TERM.

A dissertação foi estruturada em diferentes capítulos: I) uma introdução que fornece uma visão

geral dos conceitos básicos de sistemas de MPs, MPs poliméricas com base na sua formulação,

mecanismos de entrega de moléculas bioactivas, e aplicações em TERM; II) a secção de Materiais

e Métodos que descreve detalhadamente os materiais utilizados para o desenvolvimento das MPs

e dos hidrogéis, bem como as respectivas metodologias de processamento e caracterização, III) o

manuscrito completo que resultou do trabalho realizado, o qual se encontra estruturado num breve

resumo, uma introdução mais concisa, o resumo dos materiais utilizados e das metodologias

seguidas ao longo do trabalho, descrição dos resultados e subsequente discussão e uma breve

conclusão; IV) as conclusões finais do trabalho e perspectivas futuras do sistema desenvolvido na

área de TERM.

Palavras-chave: Sistemas de libertação de fármacos, Engenharia de Tecidos, poli(hidroxibutirato-

co-hidroxivalerato), micropartículas, hidrogel injectável

vii

Abstract

Design and synthesis of efficient drug delivery systems (DDS) are of vital importance for tissue

engineering and regenerative medicine (TERM). Optimized delivery structures and well-defined

release kinetics appear to be logical prerequisites for safe and efficacious clinical application

avoiding undesired side effects and toxicity. With a view toward developing DDS to be applied

using minimally-invasive approaches, and that would provide sustained, prolonged and localized

release of biochemical cues relevant for tissue regeneration, we designed a biphasic injectable

DDS combining polyhydroxybutyrate-co-hydroxyvalerate (PHBV) microparticles (MPs) within a

gellan gum (GG) hydrogel.

The production of PHBV MPs incorporating proteins and glucocorticoids model molecules was the

first step of the work presented in this dissertation. For this purpose, Bovine Serum Albumin (BSA)

and Dexamethasone (Dex) were incorporated inside PHBV MPs by using a double emulsification-

solvent evaporation method with modifications. Aiming at tailoring the bioactive molecules release

profile, different proportions of chloroform and ethanol, as organic phase, were proposed. This

modification revealed to affect some features of the microparticulate system, namely particle size

and surface topography. The influence of the presence of ethanol in the organic phase was also

evaluated in terms of bioactive molecules incorporation efficiency. The presence of ethanol led to

higher efficiency of Dex entrapment inside the MPs, while the BSA incorporation was not affected.

In his turn, the in vitro release studies revealed that MPs produced in the absence of ethanol in the

organic phase presented a more sustained profile when compared to the MPs produced in the

presence of ethanol. This finding was expected since the addition of ethanol to the organic phase

induced the formation of larger pores on the surface and in the core of the MPs thus facilitating

water penetration, and consequently faster release of the bioactive molecule.

After obtaining the PHBV MPs loaded with the molecules of interest, the objective was to define a

system that would act as carrier of the MPs and would enhance MPs residence time upon injection.

For this purpose, the different PHBV MPs were embedded within injectable GG hydrogels,

extensively studied in our Research Group as cell carriers for different tissue engineering

applications. A uniform distribution of PHBV MPs across the GG matrix with a strong integration of

the MPs with the polymer of the hydrogel was attained. As expected, the incorporation of the MPs

into the GG hydrogels strongly influenced the profile, from one to nearly zero-order and the amount

of released Dex and BSA. Although the properties of the hydrogel are prone to further improvement

in order to tune the release profile according to tissue and pathology/injury specific requirements,

the proposed system can certainly be explored as a long-term DDS.

In this sense, a biphasic DDS, which allows the release of biochemical cues with different

physicochemical features, as well as its localized deliver, was successfully developed and

represents a versatile tool to prepare instructive cell microenvironments for TERM.

viii

The dissertation was structured in different chapters: I) a General Introduction that provides an

overview of the basics of microparticulate systems, polymeric microparticles based on their

formulation, their mechanisms of drug delivery, and their applications in the TERM; II) a Materials

and Methods section that describes with some detail the materials used to produce the MPs and

the hydrogels as well as the respective processing and characterization methodologies; III) the

complete manuscript that resulted from the work developed in which is provided a short abstract, a

more concise introduction of the work, a resume of the materials used and methodologies followed

along the work, the description of the results and subsequent discussion; IV) the final conclusions

of the work and future perspectives for the area.

Keywords: Drug Delivery System, Tissue Engineering, polyhydroxybutyrate-co-hydroxyvalerate,

microparticles, injectable hydrogel

ix

Table of Contents

I. General Introduction ........................................................................................... 1

1. Drug Delivery Systems ...................................................................................... 1

2. Microparticles as Drug Delivery Systems ........................................................ 3

2.1. Microparticulate Systems in Tissue Engineering .......................................... 4

2.1.1. Microparticulate Drug Delivery Systems ......................................................... 5

2.1.2. Microparticulate Drug Delivery Systems as part of Tissue Engineering Scaffolds 7

2.1.3. Microparticulate Drug Delivery Systems as Tissue Engineering Scaffolds .......... 8

3. Final Remarks and Future Trends .................................................................... 9

4. Bibliography..................................................................................................... 10

II. Materials and Methods ................................................................................... 17

1. Poly (hydroxybutyrate-co-hydroxyvalerate) Properties ................................ 17

2. Gellan Gum Properties..................................................................................... 17

3. Preparation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles . 18

3.1. Loading Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles with

bioactive molecules ......................................................................................... 18

3.1.1. Bovine Serum Albumin-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles .................................................................................................... 18

3.1.2. Dexamethasone-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles…………………………………………………………………………..19

4. Gellan Gum hydrogel preparation ................................................................... 19

4.1. Incorporation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles in

Gellan Gum hydrogel ....................................................................................... 19

5. Determination of Particles Production Yield .................................................. 19

6. Efficiency of incorporation of bioactive molecules into the Poly

(hydroxybutyrate-co-hydroxyvalerate) microparticles ............................... 20

6.1. Bovine Serum Albumin Quantification ........................................................ 20

6.2. Dexamethasone Quantification................................................................... 20

7. In vitro Release Studies ................................................................................... 21

7.1. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles .................................................................................................. 21

7.2. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles embedded within Gellan Gum hydrogels ............................................ 21

x

8. Laser Diffraction Spectrometry ....................................................................... 21

9. Fourier Transform Infrared spectroscopy ...................................................... 22

10. Scanning Electron Microscopy ..................................................................... 22

11. References ...................................................................................................... 22

III. Development of an Injectable PHBV microparticles-GG Hydrogel Hybrid

System for Tissue Engineering Applications ................................................... 25

1. Introduction ..................................................................................................... 26

2. Materials and Methods ..................................................................................... 29

2.1. Preparation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles ... 29

2.1.1. Bovine Serum Albumin-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles .................................................................................................... 29

2.1.2. Dexamethasone-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles…. ................................................................................................ 29

2.2. Gellan Gum hydrogel preparation ............................................................... 29

2.2.1. Incorporation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles in

Gellan Gum hydrogel.......................................................................................... 30

2.3. Determination of Particles Production Yield ............................................... 30

2.4. Efficiency of incorporation of bioactive molecules into the Poly

(hydroxybutyrate-co-hydroxyvalerate) microparticles ....................................... 30

2.4.1. Bovine Serum Albumin Quantification .......................................................... 30

2.4.2. Dexamethasone Quantification ................................................................... 31

2.5. In vitro Release Studies ............................................................................. 31

2.5.1. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles .................................................................................................... 31

2.5.2. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles embedded within Gellan Gum hydrogels .......................................... 31

2.6. Laser Diffraction Spectrometry ................................................................... 32

2.7. Fourier Transform Infrared spectroscopy ................................................... 32

2.8. Scanning Electron Microscopy ................................................................... 32

3. Results .............................................................................................................. 33

3.1. Poly (hydroxybutyrate–co–hydroxyvalerate) microparticles Process

Production ....................................................................................................... 33

3.2. Bovine Serum Albumin and Dexamethasone Entrapment Efficiency and

Release Profile ................................................................................................. 36

3.2.1. Loaded Microparticles properties ................................................................ 36

3.2.2. Bovine Serum Albumin Release from Poly (hydroxybutyrate–co–hydroxyvalerate)

microparticles .................................................................................................... 39

xi

3.2.3. Dexamethasone Release from Poly (hydroxybutyrate–co–hydroxyvalerate)

microparticles .................................................................................................... 41

3.3. Injectable Gellan Gum/ Poly (hydroxybutyrate–co–hydroxyvalerate)

microparticles System ...................................................................................... 43

3.3.1. Injectable System Properties ...................................................................... 43

3.3.2. Bovine Serum Albumin Release from the Injectable Gellan Gum/ Poly

(hydroxybutyrate–co–hydroxyvalerate) microparticles System ................................. 44

3.3.3. Dexamethasone Release from the Injectable Gellan Gum/Poly (hydroxybutyrate–

co–hydroxyvalerate) microparticles System ........................................................... 45

4. Discussion ........................................................................................................ 46

5. Conclusion ....................................................................................................... 50

6. References ....................................................................................................... 51

IV. Final Remarks and Future Perspectives ...................................................... 59

V. Annex ............................................................................................................... 61

1. Morphological Analysis .................................................................................. 61

xii

List of Figures

I. General Introduction

Figure 1 – Requisites regarding the design of a new drug delivery system…………………………..2

III. Development of an Injectable PHBV microspheres-GG Hydrogel Platform

Targeting Tissue Engineering Applications

Figure 1 – Schematic depiction of double emulsification evaporation method……………………..33

Figure 2 – Particle size distribution of the PHBV MPs…………………………………………………34

Figure 3 - SEM micrographs of PHBV MPs obtained under different experimental conditions: (A)

using 100% chloroform; and (B) a mixture of 90% chloroform: 10% ethanol. The sequential

micrographs represent an overview of the obtained MPs showing the detail of their surface and

respective close up of the internal morphology…………………………………………………………35

Figure 4 - SEM micrographs of BSA loaded-PHBV MPs obtained under different experimental

conditions: (A) using 100% chloroform; and (B) 90% chloroform: 10% ethanol. The sequential

micrographs represent an overview of the obtained MPs showing the detail of their surface……..37

Figure 5 – FTIR spectra of BSA and PHBV MPs: BSA (black line); 100 unloaded-PHBV MPs (red

line); 90 unloaded-PHBV MPs (blue line); 100 BSA-loaded PHBV MPs (pink line); 90 BSA-loaded

PHBV MPs (green line). The characteristics bands of BSA are marked (*)………………………….38

Figure 6 – FTIR spectra of Dex and PHBV MPs: Dex (black line); 100 unloaded-PHBV MPs (red

line); 90 unloaded-PHBV MPs (blue line); 100 Dex-loaded PHBV MPs (pink line); 90 Dex-loaded

PHBV MPs (green line). The characteristics bands of Dex are marked (*)…………………………..39

Figure 7 - In vitro release profile of BSA from BSA-loaded PHBV MPs obtained using a mixture of

chloroform and ethanol in different proportions: (100:0) % and (90:10) %.......................................40

Figure 8 – FTIR spectra of BSA-loaded PHBV MPs: BSA-loaded 100 PHBV MPs (black line) and

BSA-loaded 90 PHBV MPs (red line) before release studies; and BSA-loaded 100 MPs (blue line)

and BSA-loaded 90 MPs (purple line) after 21 days of in vitro release. The characteristics bands of

BSA are marked (*)…………………………………………………………………………………….......40

xiii

Figure 9 - SEM micrographs of BSA-loaded PHBV MPs obtained under different experimental

conditions: (A) using 100 % chloroform; and (B) with 90% chloroform: 10% ethanol, after 21 days of

in vitro release studies. The sequential micrographs represent an overview of the obtained MPs

showing the detail of their surface………………………………………………………………………...41

Figure 10 - In vitro release profile of Dex loaded-MPs obtained using a mixture of chloroform and

ethanol in different proportions: (100:0) %; and (90:10) %...............................................................42

Figure 11 – FTIR spectra of Dex-loaded PHBV MPs: Dex-loaded 100 PHBV MPs (black line) and

Dex-loaded 90 PHBV MPs (red line) before release studies; and 100 Dex-loaded MPs (line) and 90

Dex-loaded MPs (purple line) after 21 days of in vitro release. The characteristics bands of Dex are

marked (*)……………………………………………………………………………………………………42

Figure 12 – Schematic representation of injectable fluorescein isothiocyanate labelled bovine

serum albumin – loaded PHBV MPs embedded within GG hydrogel system………………………..43

Figure 13 - SEM micrographs of PHBV MPs embedded within injectable GG hydrogels. The

sequential images represent an overview of the distribution of the particles within the hydrogel,

close up on the surface of the biphasic structure, and a cross-section………………………………43

Figure 14 - In vitro release profile of BSA from GG hydrogels embedding PHBV MPs obtained

using a mixture of chloroform and ethanol in different proportions: (100:0) %; and (90:10) %........44

Figure 15 - SEM photographs of BSA loaded-MPs embedded within GG hydrogel, after 21 days of

in vitro release studies. The sequential images represent an overview of the distribution of the

particles within the hydrogel, close up on the surface of the biphasic structure, and a cross-

section……………………………………………………………………………………………………….45

Figure 16 - In vitro release of Dex from GG hydrogels embedding PHBV MPs obtained using a

mixture of chloroform and ethanol in different proportions: (100:0) %; and (90:10) %.....................45

V. Annex

Figure 1 - SEM micrographs of Dex loaded-PHBV MPs obtained under different experimental

conditions: (A) using 100% chloroform; and (B) 90% chloroform: 10% ethanol. The sequential

micrographs represent an overview of the obtained MPs showing the detail of their surface……..61

xiv

List of Tables

III. Development of an Injectable PHBV microspheres-GG Hydrogel Platform

Targeting Tissue Engineering Applications

Table I. Yield and Size of the PHBV MPs produced under different conditions……………....33

Table II. Yield and Size of the PHBV loaded MPs produced under different conditions……..36

Table III. Incorporation Efficiency of Bovine Serum Albumin and Dexamethasone within PHBV

MPs………………………………………………………………………………………………….....38

List of Abbreviations

- weight of MPs

– weight of the sum of polymer and

bioactive substance

100 PHBV MPs – Microparticles

obtained in the absence of ethanol

90 PHBV MPs - Microparticles obtained

in the presence of ethanol

BMPs - Bone Morphogenetic Proteins

BSA - Bovine Serum Albumin

DDS - Drug Delivery System

Dex - Dexamethasone

EC - Endothelial Cells

ECM - Extracellular Matrix

FTIR - Fourier Transform Infrared

GFs - Growth Factors

GG - Gellan Gum

IE - Incorporation Efficiency

IGF-1 - Insulin-like Growth Factor-1

MPs - Microparticles

Mw - Molecular Weight

O – Organic phase

OPF - Oligo(poly(ethylene glycol)

Fumarate

PGA - Poly(Glycolic acid)

PHAs - Polyhydroxyalkanoates

PHB – Poly (hydroxybutyrate)

PHBV – Poly (hydroxybutyrate-co-

hydroxyvalerate)

PLA – Poly (lactic acid)

PLGA – Poly (lactic-co-glycolic) acid

PVA - Poly(vinyl alcohol)

RM - Regenerative Medicine

SEM - Scanning Electron Microscopy

TE – Tissue Engineering

xv

TGF - Transform Growth Factor

v/v % - Volume concentration

VEGF - Vascular Endothelial Growth

Factor

w/v % - Mass concentration

W1 – First aqueous phase

W2 – Second aqueous phase

W/O/W – Water/Oil /Water

1

I. General Introduction

1. Drug Delivery Systems

Drug Delivery is a field of vital importance for medicine and healthcare (Zhang, 2013).

This field of pharmaceutical technology represents one of the most rapidly advancing

areas of science in which chemists and chemical engineers are contributing to human

health care (Jain, 2008).

Therapeutic efficacy and safety of drugs, administered by conventional methods, can be

improved by more precise spatial and temporal placement within the organism, thereby

reducing both the size and number of doses by using a controlled drug delivery system

(DDS) (Bassyouni, 2013). Indeed, the DDS plays a vital role in controlling the

pharmacological effect of the drug as it can influence its pharmacokinetic profile, release

rate, the site and duration of its action and subsequently the side-effect profile (Perrie,

2010).

Controlled DDS improves bioavailability by preventing premature degradation and by

enhancing uptake, maintains drug concentration within the therapeutic window by

controlling the drug release rate, and reduces side effects by targeting diseased site and

specific cells (Zhang, 2013). The release of the bioactive molecule is extended along with

the time which is highly beneficial for molecules that are rapidly metabolized and

eliminated from the body after administration (Jain, 2008). With this kind of release

systems, the rate of drug release matches the rate of drug clearance and, therefore, the

drug concentration is within the therapeutic window for the vast majority of the 24 hours

period.

Tissue engineering (TE) was defined by Langer and Vacanti in 1993 as “an

interdisciplinary field of research that applies the principles of engineering and life

sciences towards the development of biological substitutes that restore, maintain, or

improve tissue function”. In this research domain, the general strategy consists in

developing biodegradable/biocompatible scaffold materials that play an important role as

artificial extracellular matrix (ECM) in which cells are seeded/cultured, under pre-

determined conditions, with defined biochemical and mechanical cues (Hutmacher, 2000).

Nevertheless, they are often unable to create the exact/correct microenvironment during

the engineered tissue development to promote the accurate in vitro tissue development.

The emerging and promising next generation of engineered tissues is relying on

producing scaffolds with functional cues (Malafaya, 2007) aiming at driving host healing

2

response at the site of injury. In order to produce biofunctional engineered tissues,

researchers have been combining scaffolds with cells from assorted sources and/or

growth factors (GFs) to stimulate cell migration, differentiation and tissue remodelling

(Xiao, 2003; Kanczler, 2008; Anderson, 2009; Chen, 2010). With an improved

understanding of the critical pathways involved in the development of integrated tissues,

the role of GFs in tissue regeneration, and the expansion of their availability through

recombinant technologies, the delivery of GFs by DDS is an increasingly important

strategy to repair or regenerate damaged/diseased tissue and is a leading component of

tissue engineering approaches.

In the context of tissue engineering, an ideal DDS must have several general

requirements, including being a biodegradable carrier material, possess high loading

efficiency, a controlled release profile, the ability to target or be retained at the desired site

of action, and ideally a certain degree of versatility that for example allows several

bioactive molecules to be sequentially released in a manner that mimics the temporal

profile of the healing process in vivo (Vasita, 2006; Chen, 2009a; Balasubramanian, 2010).

Simultaneously, several factors, such as the properties of the drug/bioactive molecule,

route of administration, nature of delivery vehicle, mechanism of drug release, ability of

targeting, and biocompatibility must be taken in consideration (Figure 1) (Chiellini, 2008).

Figure 1 – Requisites regarding the design of a new drug delivery system (adapted from Chiellini,

2008).

Drug Properties

Route of Administration

Biocompatibility

Ability of Targeting

Mechanism of Drug Release

Delivery System Properties

Controlled Drug

Delivery Systems

3

The use of molecular or macromolecular entities and derived superstructures concerning

the delivery of drugs has a long history. Antibodies, for instance, were suggested early in

the last century as a mean to direct anticancer drugs to tumour cells in the body

expressing the corresponding antigen (Ritz, 1982). Their use as monoclonal molecules is

now in the vanguard of targeted therapy. Following advances in the discovery of cell

receptors, receptor-binding macromolecules were added to the armamentarium of DDS

for the targeting of drugs. In parallel, since the early 1970s, liposomes are at the forefront

of drug and vaccine design owing to their well-documented abilities to act as delivery

vehicles (Gregoriadis, 1993; Henriksen-Lacey, 2011). These superstructures, formed

spontaneously from amphipathic lipid molecules, together with a diverse collection of

other promising superstructures derived from a huge variety of natural and synthetic

monomeric or polymeric units, have evolved to sophisticated versions of DDS. The

incorporation onto their surface of macromolecules that contribute to optimal

pharmacokinetics of actives and their delivery to where they are needed has been

successively proposed (Vemuri, 1995; Fahy, 2009).

Recently, the advances in the peptide, proteins and drug designs and the emergence of

gene therapy have been intensifying the need for the development of new controlled

release strategies (Jain, 2008). Their clinical success is considered to be dependent on

the controlled release devices design, which must ensure that the drug release will

perfectly reach the targeted cells at a specific time. A number of DDS are currently under

investigation to circumvent the limitations commonly found in conventional dosage forms

and to improve the potential of the respective drugs. These DDS have been envisioned as

liposomes (Drulis-Kawa, 2010; Monteiro, 2013), micelles (Wilson, 2013; Xu, 2013),

polymeric micro and nanoparticles (Joshi, 2013; Liang, 2013; Kozielski, 2013; Wang,

2013), cyclodextrins (Chen, 2011; Dhule, 2012), among others (Oliveira, 2011a).

2. Microparticles as Drug Delivery Systems

There are various approaches to deliver a therapeutic substance to the target site in a

sustained controlled release fashion. Of the different forms reported, microparticles

attained much importance due to a tendency to accumulate in inflamed areas of the body

(Puddu, 2010). The terminology used to describe microparticulate formulations can

sometimes be inconsistent and confusing to readers unfamiliar with the field. Basically,

the term “microparticle” refers to a particle with a diameter of 1-1000 μm, irrespective of

the precise internal or external structure (Arshady, 1988; Freiberg, 2004). However, this

criterion is very ubiquitous, and some researchers have been considering that, in terms of

4

diameter, microparticles cover a range between 0.1 and 2 μm (Shet, 2008; Puddu, 2010;

Siljander, 2011, Press, 2012). Within the broad category of microparticles, microspheres

specifically refers to core-shell microparticles, whilst the subcategory of microcapsules

applies to microparticles which have a core surrounded by a material, distinctly different

from that of the core, which can be solid, liquid, or even gas (Chemtob, 1986; Brazel,

2000; Berkland, 2004).

A number of different polymers, both synthetic and natural, have been exploited for the

fabrication of biodegradable microparticles (Oliveira, 2011b). Synthetic polymers have the

advantage of sustaining the release of the encapsulated therapeutic agent over a period

of days to several weeks. In opposition, natural polymers have a relatively short duration

of drug release and are in general limited by the use of organic solvents and relatively

harsher processing conditions (Panyam, 2003).

The most frequently used polymers for this purpose include poly(lactic acid) (PLA),

poly(glycolic acid) (PGA), their copolymers poly(lactic-co-glycolic) acid (PLGA), poly-ε-

caprolactone, polyethylene, polymethyl methacrylate, oligo(poly(ethylene glycol) fumarate)

(OPF), microparticulate systems; natural polymers such as poly(hydroxybutyrate) (PHB),

its copolymers poly(hydroxybutyrate-co-hydroxyvalerate) (PHBV), alginate or gelatin

(Vasita, 2006; Balasubramanian, 2010). In addition, inorganic materials such as low and

high temperature calcium orthophosphates (including calcium phosphate cements and

sintered ceramics), calcium sulphate cements and Bioglass have also been used as GF

carriers (Chen, 2009b).

2.1. Microparticulate Systems in Tissue Engineering

Controlled release has many direct applications to TE. For example, local delivery of GFs

can be accomplished by encapsulating the agent within a biocompatible polymer matrix or

microparticle. The controlled-release polymer system is then implanted at the desired

tissue site, where it releases the soluble factor directly into the interstitial space of the

tissue (Saltzman, 2002).

GFs can be incorporated into the scaffolds by two approaches. The first approach

involves adding a lyophilized factor like vascular endothelial growth factor (VEGF) to the

polymeric microparticles prior to processing the polymer into a porous scaffold, which

results in the factor being largely associated with the surface of the polymer. In this

method, VEGF is subjected to rapid release (e.g., days to weeks in duration). The second

approach involves pre-incorporation of a GF in microspheres. Therefore, fabricating

scaffolds from these particles results in a more even distribution of factors throughout the

5

polymer, with release kinetics controlled by degradation rate of the polymer used to

construct microparticles. The two approaches may be combined by mixing particulate

polymers containing the first factor with microspheres containing a pre-encapsulated

second factor to deliver two GFs with different release rates. The particulate and

microparticles are then fused to form a homogeneous combined scaffold with an open-

pore structure (Richardson, 2001).

2.1.1. Microparticulate Drug Delivery Systems

Therapeutic induction of tissue repair or function restoration in tissue engineering and

regenerative medicine requires recapitulation of the spatial and temporal

microenvironments presented by natural ECM, in many cases by providing exogenous

instructive bioactive signals (Biondi, 2008). The identification and production of

recombinant morphogens and GFs that play key roles in tissue regeneration have

generated much enthusiasm. Pivotal among these signals are GFs, a group of proteins

capable of acting on cell-surface receptors and directing cellular activities involved in

wound healing, such as bone morphogenetic proteins (BMPs, particularly BMP-2, -4 and -

7), VEGF, epidermal growth factor, and fibroblast growth factor, which provide vital

stimulation of cell recruitment, proliferation, morphogenesis, and differentiation (Tabata,

2000; Chen, 2010).

To be effective as a therapeutic agent, a GF has to reach the site of injury without

degradation, and then, it has to remain in the target location sufficiently long to exert its

action(s) (Chen, 2009b). GFs that are provided exogenously in solution into the site to be

regenerated are generally not effective because GFs tend to diffuse away from injury sites

and to be enzymatically digested or deactivated (Chen, 2009b; Chen, 2010). Interestingly,

clinical trials of DDS that have shown benefit all contain a common denominator, the

presence of a material carrier, strongly suggesting that spatiotemporal control over the

location and bioactivity of factors after introduction into the body might be achieved by the

use of those carriers (Lee, 2011).

In the search for vehicles that delivery without being damaged, microparticulate systems

made of biodegradable polymers as reveal significant advantages when incorporating

GFs. These systems have the potential to provide sustained release kinetics in remote

parts of the body after implantation (Chen, 2010). The release profile of a GF can be

altered either by tailoring the properties of the polymer or polymeric composition or even

by adjusting the physical and chemical properties of the microparticles such as surface

porosity, particle size, and its degradation rate (Chen, 2007; Chen, 2009b). Moreover, the

6

drug loading capacity of particulate systems can be greatly improved if a capsule is

formed or if the core of each microsphere is designed to have an interconnected pore

structure. Microparticles are not internalized by the cells and are retained in the tissue,

providing prolonged bioactive substance release. Additionally, it is not easy for

microparticles to diffuse out of the target tissue to other sites to cause deleterious side-

effects in other tissues because of their size. Thus, they provide good control over the

release rate and dose of GFs in a local microenvironment, yielding desirable

concentrations over a period of days or months (Varde, 2004; Anitua, 2008; Mundargi,

2008).

Microparticles carrying GFs, small molecules, among others have been prepared by

various techniques and their role has been investigated both in in vitro and in vivo. PLGA -

based microcapsules containing GFs have been prepared by a double emulsion–solvent

evaporation technique (Wenk, 2009). In 2008, Rocha and co-workers showed that the

sustained delivery of VEGF encapsulated in PLGA microspheres interferes with wound

healing cascade events and its release upregulates angiogenesis (Rocha, 2008). The

degree of encapsulation and mechanism of growth factor entrapment are dependent on

hydrophobic–hydrophobic or hydrophilic–hydrophilic interactions among the molecules

and polymers. Blending of materials allow more complex structures and patterns of factor

entrapment (Lee, 2011). In 2009, Zhu and collaborators have proposed a blend of PLGA

and PHBV microparticles prepared by an emulsion technique to release of hepatocyte

growth factor (Zhu, 2009). The final microparticulate structure showed a core-shell

structure, where PHBV molecules, which are more hydrophobic and less degradable,

were distributed within the shell. This prevents the loss of GFs during processing,

resulting in a better encapsulation material than microspheres of PLGA- or PHBV-alone

(Zhu, 2009). Recent studies with calcium phosphate cement /gelatin composite and

diopside (CaMgSi2O6) ceramic microspheres showed that ceramic microparticle systems

can be successfully used as injectable, biodegradable and osteoinductive delivery

vehicles (Leeuwenburgh, 2010). Kirby and co-workers also demonstrated that

PLGA/PLGA-PEG-PLGA microparticles entrapping and releasing BMP-2 in a sustained

manner promote the differentiation of MC3T3-E1 cells to a greater extent than osteogenic

supplements (Kirby, 2011).

7

2.1.2. Microparticulate Drug Delivery Systems as part of Tissue Engineering

Scaffolds

When microparticles are injected in target tissues, they are continuously exposed to

external mechanical deformation, leading to the uncontrollable displacements of particles

(Fitzgerald, 1987; Griffith, 2000; Chan, 2005). This displacement greatly affects the

concentration of bioactive macromolecules in target tissues, resulting in limited tissue

regeneration (Lemperle, 2004). Moreover, when in free movement the particulate system

can be expelled even before release of the therapeutic agent. Furthermore, these

approaches for therapeutic tissue regeneration usually involve one-time delivery of single

biochemical cues. Therefore, extensive efforts are being made to tackle these limitations

while harnessing the advantages of particulate systems, including their minimally invasive

delivery and controlled drug release.

The development of sophisticated drug delivery systems with multiple functionalities is

one of the challenges of TE. Extraordinary progress has been made in the last years

regarding the design of scaffolds with suitable multiscale hierarchical structure and toward

the design of delivery systems able to release active molecules following a hypothetically

complex delivery pattern (Biondi, 2008).

Current TE strategies are focusing on the development of porous scaffolds specifically

designed to mimick tissues ECM at morphological but most importantly at biochemical

level by incorporating bioactive components, selected to promote full regeneration

(Guldberg, 2009). In this sense, the concept of particulate systems further associated with

3D scaffolds (Chen, 2007; Chung, 2007; Fei, 2008) serves two functions, as a scaffolding

structure for cell attachment and tissue ingrowth, as well as a delivery platform for multiple

bioactive molecules release. A variety of systems have been designed with this purpose.

Chen and co-workers proposed a dual release system of BMP-2 and Insulin-like Growth

Factor-1 (IGF-1) from gelatin microparticles embedded in a glycidyl methacrylated dextran

hydrogel for periodontal TE purposes (Chen, 2009b). In this study, it was shown that the

system allowed the release of both proteins in a sustained and independent fashion,

which facilitated cell attachment, proliferation and osteoblastic differentiation of

periodontal ligament fibroblasts in a synergistic manner. The effect of simultaneous,

controlled release of VEGF and monocyte chemotactic protein-1 from dual particulate

systems as a strategy for therapeutic vascularization was proposed by Jay (2010) and

collaborators. Alginate microparticles loaded with both VEGF and MCP-1 with distinct

release kinetics were integrated into a collagen/fibronectin gel used as endothelial cells

(EC) carrier. The combined delivery of VEGF and MCP-1 increased functional vessel

formation from transplanted EC and also led to a higher number of smooth muscle cell-

8

invested vessels than did EC therapy alone. Despite the well-known role of MCP-1 in

inflammation, these beneficial effects were accomplished without a long-term increase in

monocyte/macrophage recruitment or a shift to a pro-inflammatory (M1) macrophage

phenotype. More recently, stem cells encapsulated in OPF hydrogels combined with

transform growth factor (TGF)-β3 and/or IGF-1-loaded gelatin MPs were proposed for

osteochondral tissue regeneration by Kim et al. (2013). In this study, a bilayered OPF

hydrogel that intended to mimic the distinctive hierarchical structure of native

osteochondral tissue was used to assess the effect of TGF-β3 with varied release kinetics

on osteochondral tissue regeneration in a rabbit full-thickness osteochondral defect

model. The in vivo study revealed that after 12 weeks post-implantation IGF-1 delivery

alone contributed to enhanced cartilage repair in comparison to the dual GFs delivery.

Additionally, no significant effects of the TGF-β3 release kinetics were observed on

osteochondral tissue repair. The results suggested that the dual delivery of TGF-β3 and

IGF-1 may not synergistically enhance the quality of engineered tissue, revealing that

combined and complex approaches do not necessarily confer improved healing over the

single delivery of GFs, and that most likely the approach to follow is highly dependent on

the TE application. Thus, although current evidences suggest that well-chosen GFs

cocktails can represent an advantage when associated in simultaneous delivery systems,

uncertainties whether reciprocal or cooperative interactions exist as fare to be clear.

2.1.3. Microparticulate Drug Delivery Systems as Tissue Engineering Scaffolds

The use of microparticles as scaffolds is a relatively poorly explored TE approaches that

has many advantages, mainly due to its versatility (Borden, 2002; Lu, 2003; Zhang, 2003;

Ungaro, 2006). In fact, microspheres can be easily modified in a controlled manner (Hong,

2005) to introduce various ECM proteins aiming to tailor the cell-material interaction

promoting cell adhesion. Simultaneously, GFs as well as other bioactive molecules can be

easily incorporated and then further released in a controlled manner from the

microspheres that acted as scaffolding structures themselves, thus regulating cell

behavior (Tabata, 1999; Perets, 2003; Chen, 2003). The so-called microparticles

aggregation method consists in their aggregation, by physical or chemical means, forming

those 3D structures (Maquet, 1997) that can provide support for cell adhesion and also

act as carriers of bioactive molecules (Gomes, 2005). For microsphere-based scaffolds,

microsphere size is one of the major determinants of polymer degradation rate, governing

the release kinetics of loaded molecules and providing the control over pore sizes and

macro-porosity (Singh, 2008).

9

A different approach was developed by Nof and co-workers (Nof, 2002) which produced

PLGA microspheres by double emulsification process incorporating a DNA plasmid and

further designed using a gas foaming technique to create the 3D porous structures of

different shapes. The obtained scaffolds exhibited a sustained plasmid release for at least

21 days with minimal burst effect during the initial phase, which differed from the first 24

hours release profile from individual microparticles. In a study by Malafaya et al., a novel

approach was used envisioning osteochondral regeneration by effective differentiation of

adipose tissue-derived mesenchymal stem cells in osteogenic and chondrogenic media.

In this study, scaffolds were obtained by the agglomeration of chitosan microparticles

using a heat-induced process. Osteochondral bilayered scaffolds were also developed

consisting of a hydroxyapatite part and another made up of chitosan, linked by chemical

crosslinking process, leading to an integrative bone and cartilage interface (Malafaya,

2005).

3. Final Remarks and Future Trends

Successful repair and regeneration strategies will require quantitative insight of tissue

microenvironment and can be engineered via designing biomaterials, which provide

quantitative adhesion, growth, or migration signals to direct cellular differentiation

pathways. The use of microparticles in the health sciences has been adapted to TE

strategies in several approaches. Advances in bioactive materials as well as DDS allow

not only controlled release, but also protection of bioactive molecules from degradation.

The combination of microparticles with TE scaffolds allow the control over local

concentrations needed and/or local concentration gradients needed for successful tissue

regeneration. Moreover these systems can be processed into a scaffold constituting a

versatile platform for cell adhesion and bioactive molecules release.

From an upstream perspective, an effort is being made to develop scalable differentiation

technologies and to study morphogen gradients in pluripotent stem cell differentiation by

using localized delivery of GFs within multicellular aggregates from microparticle delivery

vehicles. For this purpose, some advances are being made and researchers have

reported gelatin-based microparticles to deliver GFs to specific areas of embryoid bodies

as aggregates of differentiating stem cells. Although huge developments have being

made, these release technologies in TE&RM is not yet delivering significant progress in

terms of clinical outcomes and commercialization, this fascinating field of research is

10

bound to dramatically change clinical practice and the therapeutic choices made by

clinicians, resulting in significant therapeutic commercial devices and in vitro models.

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17

II. Materials and Methods

1. Poly (hydroxybutyrate-co-hydroxyvalerate) Properties

Polyhydroxyalkanoates (PHAs) are produced by a wide variety of microorganisms as an

internal carbon and energy storage, as part of their survival mechanism (Doi, 2002; Jung,

2005). Poly (hydroxybutyrate-co-hydroxyvalerate) (PHBV) is a member of PHAs, which

together with is homopolymer poly(hydroxybutyrate) (PHB) are the most widely used in

Tissue Engineering and Drug Delivery Systems. In terms of physicochemical features,

PHBV is thermoplastic optically active polyester which is insoluble in water and exhibit a

high degree of polymerization (Scandola, 1997; Kang, 2001). These polymers also

present some interesting characteristics, namely antioxidant (i.e. inhibits the oxidation of

other molecules) and piezoelectric properties, i.e. the capacity of a material to suffer

electric polarization due to mechanical stress (Chen, 2000). PHB and PHBV, which are

members of PHAs family, degrade into d-3-hydroxybutyrate acid, which is a normal

constituent of human blood, which may justify their low toxicity (Tokiwa, 2004; Choi, 2005;

Cheng, 2006). PHB is very crystalline and brittle, whereas the copolymers of PHB with

hydroxyvaleric acid (PHBV) are less crystalline (Srithep, 2013), more flexible, and more

processable (Barham, 1984; Yu, 2006).

2. Gellan Gum Properties

Gellan gum (GG) is one of the widely used fermentation materials, which offers a solution

to many problems encountered in gelling agents, which is resistant to heat and acid

(Prajapati, 2013). GG is an extracellular microbial anionic heteropolysaccharide consisting

of glucose–glucuronic acid–glucose–rhamnose as a repeating unit and that forms a gel in

the presence of metallic ions (Kang, 1982; Jansson, 1983). It is commercially available in

two forms, acetylated and deacetylated both forming thermo-reversible gels with different

mechanical properties in the presence of metallic ions and upon temperature decrease.

The key advantages of GG are its high gelling efficiency and its ability to produce a wide

spectrum of mechanical properties (Sworn, 2009). GG also requires concentrations two to

three times less than those of gelatin and agar to achieve the similar levels of mechanical

strengths (Morris, 2012). Moreover, GG based hydrogels have been shown to efficiently

sustain the deliver and growth of human cells for different tissue engineering applications

(Silva-Correia, 2012).

18

3. Preparation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles

The PHBV (molecular weight (Mw) = 425,692 g mol-1) (PHB Industrial S.A., Brazil)

microparticles were prepared by using a modification of the double emulsion solvent

evaporation method previously described by Ogawa (Ogawa, 1988). The W/O/W double

emulsion technique was followed since it offers the possibility of incorporating both

hydrophobic and hydrophilic molecules, allowing the incorporation of several agents

(Freiberg, 2004). The first step of this method is the formation of water in oil (W1/O)

emulsion where the aqueous solution (W1) can contain or not the hydrophilic active

component and the organic phase (O) contains a hydrophobic polymer. In detail, 3 ml of a

0.5% w/v Poly(vinyl alcohol) (PVA) (Sigma-Aldrich, Germany) aqueous solution (W1) was

mixed with 15 mL of a 3% w/v PHBV solution in chloroform-ethanol (O). Two different

ratios of chloroform-ethanol were used in order to verify its effect over the release profile

of the bioactive substances, namely 100:0 chloroform:ethanol (100 PHBV MPs) and 90:10

chloroform:ethanol (90 PHBV MPs). The mixture was emulsified for 3 minutes using an

ultra-turrax T18 (IKA-Werke, Germany) in order to obtain the first emulsion (W1/O). The

primary emulsion was then dripped into 150 mL of 1% w/v PVA solution (W2) and

homogenized (RW 16 basic IKA-Werke, Germany) during 5 minutes forming the

secondary emulsion (W1/O/W2). The final W1/O/W2 emulsion was left to evaporate under

magnetic stirring for 4–5 h. The obtained microspheres were collected by centrifugation,

washed at least three times with deionized water and freeze-dried.

3.1. Loading Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles with

bioactive molecules

After optimizing the procedure of preparation of PHBV MPs it was proceeded to the

incorporation of Bovine Serum Albumin (BSA) and Dexamethasone (Dex), respectively as

protein and hydrophilic bioactive substance model, and glucocorticoid as well as

hydrophobic molecule model.

3.1.1. Bovine Serum Albumin-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

BSA-loaded PHBV MPs were also prepared following the same protocol previously

described. In this sense, BSA (Sigma-Aldrich, Germany) was used at a concentration

0.1% (w/v), being dissolved within the first aqueous phase (W1).

19

3.1.2. Dexamethasone-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

A similar procedure described above was followed aiming at incorporate Dex (Sigma-

Aldrich, Germany) within the PHBV microparticles (MPs). Regarding this, Dex was

dissolved in the organic phase at a final concentration of 3.3x10-3 % (w/v).

4. Gellan Gum hydrogel preparation

For the preparation of the injectable GG hydrogels the procedure previously described by

Silva (Silva, 2013) was followed. Gelzan CM (Sigma-Aldrich, Germany) powder was

mixed at room temperature with deionized water at a concentration of 1.25% (w/v) under

constant stirring. The solution was heated at 90ºC and kept at this temperature for 30 min.

Subsequently, CaCl2 (VWR, USA) was added to the GG solution at a concentration of

0.18% (w/v) to act as a crosslinker. Finally, the solution was left at room temperature

during approximately one hour, allowing the formation of the gel, which was maintained in

a phosphate-buffered saline (PBS) solution.

4.1. Incorporation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles in

Gellan Gum hydrogel

To incorporate PHBV MPs into the GG hydrogel, a similar procedure previously described

was followed. In this regard, 30 mg of PHBV MPs were dispersed into 20 mL of a 1.25%

(w/v) GG solution. This solution was posteriorly cross-linked with 0.18% (w/v) of CaCl2

and left at room temperature during approximately one hour, being maintained in a PBS

solution.

5. Determination of Particles Production Yield

The yield of the particles prepared under the different conditions defined in section 3 was

calculated using Equation 1.

20

( ) (Equation 1)

The yield was calculated based on the weight of MPs ( ) and compared to the weight of

the compounds used to prepare the MPs ( ), polymer or the sum of polymer and

bioactive substance.

6. Efficiency of incorporation of bioactive molecules into the Poly

(hydroxybutyrate-co-hydroxyvalerate) microparticles

6.1. Bovine Serum Albumin Quantification

Aiming at determining BSA incorporation efficiency (IE) into the MPs, 10 mg of BSA-

loaded MPs were completely dissolved in chloroform with vigorous shaking and at room

temperature for 24 hours. This allows the dissolution of PHBV and at the same time

permits the release of the entrapped BSA. Then 10 mL of PBS were added in order to

dissolve the BSA. After centrifugation, the supernatant was collected and filtered through

0.2 mm membrane. Then, the concentration of incorporated protein was measured using

a Micro BCATM Protein Assay Kit (Pierce, IL) and measured at 562 nm by Synergy HT

multi-detection microplate reader (Bio-Tek's Gen5™, USA). This assay is based on two

chemical reactions. The first is the reduction of cupric ions (Cu+2) to cuprous ions (Cu+1) by

the peptide bonds, known as the biuret reaction, and by the presence of four amino acids

(cysteine, cystine, tryptophan and tyrosine) in an alkaline environment (Wiechelman,

1988; Huang, 2010). The second step is the chelation of one Cu+1 with two bicinchoninic

acid molecules, which forms an intense purple complex, which has a peak absorbance at

562nm (Smith, 1985). The protein concentration in a solution is determined by comparing

this absorbance with a standard curve of absorbance from known concentration of BSA

varying from 0 – 3.0% (w/v). The IE of BSA within PHBV MPs was calculated by equation

2, as follows:

( )

(Equation 2)

6.2. Dexamethasone Quantification

PHBV MPs with incorporated Dex were dissolved as described in section 1. Dex

concentration in the solution was measured at 242 nm by Synergy HT multi-detection

microplate reader in a 96-well quartz plate (n = 3). The amount of Dex was extrapolated

21

from a standard curve prepared with Dex solutions with known concentration in a range of

0 - 3x10-2 % (w/v). The IE of Dex was calculated as follows:

( )

(Equation 3)

7. In vitro Release Studies

7.1. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

In vitro release from PHBV MPs of both model bioactive substances was evaluated up to

21 days. Approximately, 30 mg of MPs were incubated in 10 mL of PBS (0.01 M, pH =

7.4), and maintained in a precision water bath (Grant, UK) at 60 rpm and 37±0.5ºC. At

defined time points, first at 0, 15, 30 min, and then between 1 and 8 hours, followed by 2,

3, 4, 5, 7, 14 and 21 days, 1 mL of supernatant was collected and an equivalent amount of

fresh PBS at 37 ºC was added to maintain the total volume of the sample. The

concentrations of BSA and Dex in medium were measured at 562 and 242 nm,

respectively by using the microplate reader.

7.2. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles embedded within Gellan Gum hydrogels

The in vitro release of both bioactive model substances from PHBV MPs embedded within

GG hydrogels was performed during 21 days. Approximately, 220 µL of GG hydrogel

embedding PHBV MPs were incubated in 10 mL of PBS (0.01 M, pH = 7.4), and

maintained in a precision water bath. At defined time points, 1 mL of supernatant was

collected and an equivalent amount of fresh PBS at 37ºC was added to maintain the total

volume of the sample. The concentrations of BSA and Dex in medium were measured at

562 and 242 nm, respectively by using the microplate reader.

8. Laser Diffraction Spectrometry

The particle size distribution of the prepared MPs was measured by laser diffraction

spectrometry (Coulter LS 230, Coulter Electronics, USA). The dried samples were

22

suspended in a 0.2% (v/v) of Tween 80 (Sigma-Aldrich, Germany) solution and sonicated

for 5 min with an ultra-sound probe (Sonoswiss SW 6 H, Sonoswiss® AG, Switzerland)

before measurement. The obtained homogeneous suspension was used for the particle

size distribution analysis.

9. Fourier Transform Infrared spectroscopy

Fourier Transform InfraRed Spectroscopy (FTIR) analyses, under Transmission mode,

were performed in PHBV microparticles (unloaded, loaded and after release) in order to

analyse their chemical composition. With this purpose, 1 mg of microparticles were mixed

with 40 mg of Potassium Bromide (KBr) and then processed into a disc in a manual press

(161-1100 hand press, Pike technologies, Madison, WI). FTIR-KBr spectra (IR Prestige

21, Shimadzu, Japan) were recorded at 40 scans with a resolution of 4 cm-1, at

wavelengths from 1400 to 4000 cm-1.

10. Scanning Electron Microscopy

The morphology of the PHBV MPs and dehydrated PHBV MPs embedded within GG

hydrogels was analysed before and after the in vitro release studies, by scanning electron

microscopy (SEM) using a model S360 microscope (Leica Cambridge, UK). Moreover, the

cross sections of the microparticles were performed after the embedding of this particulate

system within the GG hydrogel which were submitted to several cuts. Before being

analysed, the different structures were gold sputter coated (model SC502; Fisons

instruments, UK) for 2 minutes at 15 mA. Micrographs were recorded at 5.0 kV with

magnifications of 250, 1 000, 2 000, 8 000 and 20 000.

11. References

Barham P, Keller A, Otun E, Holmes P. Crystallization and morphology of a bacterial thermoplastic:

poly-3-hydroxybutyrate. J Mat Science 1984; 19 (9): 2781–94.

Chen G, Wu Q, Xi J, Yu H, Chan A. Microbial production of biopolyesters-polyhydroxyalkanoates.

Prog Nat Sci 2000 Nov; 10: 843–50

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Cheng S, Chen G, Leski M, Zou B, Wang Y, Wu Q. The effect of D,L-beta-hydroxybutyric acid on

cell death and proliferation in L929 cells. Biomaterials 2006 Jul; 27 (20): 3758-65.

Choi G, Kim H, Kim Y, Rhee Y. Biocompatibility of poly(3-hydroxybutyrate-co-3-hydroxyvalerate)

copolyesters produced by Alcaligenes sp MT-16. Biotechnol Bioprocess Eng 10 (6): 540–5.

Doi Y, Steinbuchel A. Biopolymers: Polyesters III - Applications and Commercial Products. Wiley-

VCH 2002 Jan; 4: 398.

Freiberg S, Zhu X. Polymer microspheres for controlled drug release. Int J Pharm 2004 Sep; 282

(1-2): 1-18.

Huang T, Long M, Huo B. Competitive Binding to Cuprous Ions of Protein and BCA in the

Bicinchoninic Acid Protein Assay. Open Biomed Eng J 2010 Nov; 4: 271-8.

Jansson P, Lindberg B, Sandford P. Structural studies of gellan gum, an extracellular

polysaccharide elaborated by Pseudomonas elodea. Carbohydr Res 1983 Dec; 124 (1):

135–9.

Jung I, Phyo K, Kim K, Park H, Kim I. Spontaneous liberation of intracellular polyhydroxybutyrate

granules in Escherichia coli. Res Microbiol 2005 Sep; 156 (8): 865-73.

Kang I, Choi S, Shin D, Yoon S. Surface modification of polyhydroxyalkanoate films and their

interaction with human fibroblasts. Int J Biol Macromol 2001 Mar; 28 (3): 205-12.

Kang K, Veeder G, Mirrasoul P, Kaneko T, Cottrell I. Agar-Like Polysaccharide Produced by a

Pseudomonas Species: Production and Basic Properties. Appl Environ Microbiol 1982

May; 43 (5): 1086–91.

Morris E, Nishinari K, Rinaudo M. Gelation of gellan—a review. Food Hydrocolloids 2012 Aug; 28

(2): 373–411.

Ogawa Y, Yamamoto M, Takada S, Okada H, Shumamoto T. Controlled-release of leuprolide

acetate from polylactic acid or copoly (lactic/glycolic) acid microcapsules: Influence of

molecular weight and copolymer ratio of polymer. Chem Pharm Bull 1988 Apr; 36 (4):

1502-7.

Prajapati V, Jani G, Zala B, Khutliwala T. An insight into the emerging exopolysaccharide gellan

gum as a novel polymer. Carbohydr Polym 2013 Apr; 93 (2): 670-8.

Scandola M, Focarete M, Adamus G, Sikorska W, Baranowska I, Świerczek S, et al. Polymer

blends of natural poly(3-hydroxybutyrate-co-3-hydroxyvalerate) and a synthetic poly(3-

hydroxybutyrate). Characterization and biodegradation studies. Macromolecules 1997 May;

30 (9): 2568–74.

24

Silva L, Cerqueira M, Sousa R, Marques A, Correlo V, Reis R. Gellan Gum‐Based Spongy‐Like

Hydrogels: Methods And Biomedical Applications Thereof (2013) Patente de Invenção

Nacional nº 106890.

Silva-Correia J, Miranda-Gonçalves V, Salgado A, Sousa N, Oliveira JM, Reis RM, et al.

Angiogenic potential of gellan-gum-based hydrogels for application in nucleus pulposus

regeneration: in vivo study. Tissue Eng Part A 2012 Jun; 18(11-12): 1203-12.

Smith P, Krohn R, Hermanson G, Mallia A, Gartner F, Provenzano M, et al. Measurement of

protein using bicinchoninic acid. Anal Biochem 1985 Oct; 150 (1): 76-85.

Srithep Y, Ellingham T, Peng J, Sabo R, Clemons C, Turng L, et al. Melt compounding of poly (3-

hydroxybutyrate-co-3-hydroxyvalerate)/nanofibrillated cellulose nanocomposites. Polym

Degrad and Stabil 2013 Aug; 98 (8): 1439-49.

Sworn G. Gellan gum. In: Phillips G, Williams P, editors. Handbook of Hydrocolloids. Cambridge:

Woodhead Publishing Ltd.; 2009. p. 204-227.

Tokiwa Y, Calabia B. Degradation of microbial polyesters. Biotechnol Lett 2004 Aug; 26 (15): 1181-

9.

Wiechelman K, Braun R, Fitzpatrick J. Investigation of the bicinchoninic acid protein assay:

identification of the groups responsible for color formation. Anal Biochem 1988 Nov; 175

(1): 231-7.

Yu L, Dean K, Li L. Polymer blends and composites from renewable resources. Prog Polym Sci

2006 Mar; 31 (6): 576–602.

25

III. Development of an Injectable PHBV microparticles-GG

Hydrogel Hybrid System for Tissue Engineering Applications

D. P. Pacheco1,2

, R. L. Reis1,2

, A. P. Marques1,2

, V. M. Correlo1,2

,

1 3Bs Research Group - Biomaterials, Biodegradables and Biomimetics, University of Minho, Headquarters of the

European Institute of Excellence on Tissue Engineering and Regenerative Medicine, AvePark, 4806-909 Taipas,

Guimarães, Portugal

2 ICVS/3B’s - PT Government Associate Laboratory, Braga/Guimarães, Portugal

Abstract

Design and synthesis of efficient drug delivery systems (DDS) are of vital importance for tissue

engineering and regenerative medicine (TERM). With a view toward developing DDS to be applied

using minimally-invasive approaches, and that would provide sustained, prolonged and localized

release of biochemical cues relevant for tissue regeneration, we designed a biphasic injectable

DDS combining Polyhydroxybutyrate-co-hydroxyvalerate (PHBV) microparticles (MPs) within a

gellan gum (GG) hydrogel.

Model molecules of proteins, Bovine Serum Albumin (BSA), and glucocorticoids, Dexamethasone

(Dex), were incorporated inside PHBV MPs by using a double emulsification-solvent evaporation

method with modifications. Aiming at tailoring the bioactive molecules release profile, different

proportions of chloroform and ethanol, as organic phase, were proposed. This modification

revealed to affect some features of the microparticulate system, namely particle size and surface

topography. The presence of ethanol led to higher efficiency of Dex entrapment inside the MPs,

from 82 to 92 %, approximately, but did not affect BSA incorporation, approximately 66%

independently of the presence of ethanol.

In his turn, the in vitro release studies revealed that MPs produced in the absence of ethanol in the

organic phase presented a more sustained profile when compared to the MPs produced in the

presence of ethanol. This finding was expected since the addition of ethanol to the organic phase

induced the formation of larger pores on the surface and in the core of the MPs thus facilitating

water penetration, and consequently faster release of the bioactive molecule.

In his turn, the in vitro release results indicated that around 68% of Dex incorporated in the MPS

produced in the absence of ethanol was released after 21 days, while from the particulate system

produced in the presence of ethanol 100% of the incorporated Dex was released after 5 days.

Regarding BSA, the presence of ethanol during MPs fabrication did not affect the amount of protein

released after 21 days of immersion, around 78% independently of the conditions. In order to

define a system that would act as carrier of the MPs enhancing MPs residence time upon injection,

the obtained PHBV MPs were embedded within injectable GG hydrogels. As expected, the

incorporation of the MPs into the GG hydrogels strongly influenced the profile, from one to nearly

zero-order and the amount of released Dex and BSA. However, a uniform distribution of PHBV

MPs across the GG matrix with a strong integration of the MPS was attained.

In this sense, a biphasic DDS, which allows the release of biochemical cues with different

physicochemical features, as well as its localized deliver, was successfully developed and

represents a versatile tool to prepare instructive cell microenvironments for TERM.

26

1. Introduction

Design and synthesis of efficient drug delivery systems (DDS) are of vital importance for

tissue engineering (TE) and regenerative medicine (RM) applications (Biondi, 2008) as a

way to maintain a particular active agent’s concentration in patients’ blood and/or tissues

for an extended time, without the need of additional administration (Amass, 1998; Kost,

2001). So far, these systems have been relying on liposomes (Drulis-Kawa, 2010;

Monteiro, 2013), micelles (Wilson, 2013; Xu, 2013), polymeric micro and nanoparticles

(Joshi, 2013; Liang, 2013; Kozielski, 2013; Wang, 2013), cyclodextrins (Chen, 2011;

Dhule, 2012), among others (Oliveira, 2011). Drug carriers for implantable or injectable

DDS should have good biocompatibility, biodegradability and controlled drug

delivery/release capability.

Biodegradable polymers, either synthetic or natural, are capable of being cleaved into

biocompatible byproducts through chemical or enzyme-catalyzed hydrolysis. This

biodegradable property makes it possible to implant them into the body with predicted

release profile and without the need of subsequent removal by surgical operation. Drugs

formulated with these polymers can be released in a controlled manner, by which the drug

concentration in the target site is maintained within the therapeutic window. The release

rates of the drugs from biodegradable polymers can be controlled by a number of factors,

such as biodegradation kinetics of the polymers (Ciçek, 1995; Zhang, 2002; Mi, 2002),

physicochemical properties of the polymers and drugs (Calandrelli, 2002; Abraham,

2003), thermodynamic compatibility between the polymers and drugs (Liu, 2004), and the

shape of the devices (Chen, 2001; Tunón, 2003; Fulzele, 2004).

Several based biodegradable polymers have been used for this purpose, including

poly(lactide-co-glycolide) (Wen, 2013), poly(l-lactic acid) (Correia, 2013), chitosan

(Champa, 2010), alginate (Iwanaga, 2013), starch-poly-ε-caprolactone (Balmayor, 2009),

poly(hydroxybutyrate) (PHB) (Shishatskaya, 2008) and poly(hydroxybutyrate-co-

hydroxyvalerate) (PHBV) (Chen, 2012).

Polyhydroxyalkanoates (PHAs) have been intensively investigated as a family of natural

biodegradable and biocompatible materials for in vivo applications as implantable TE

material, as well as, release vectors for various drugs (Kabilan, 2012). PHAs have

physicochemical properties similar to the widely used synthetic polymers (e.g. propylene

and polyethylene), which are non-biodegradable being most of the time compared with

poly(lactide-co-glycolide), a biodegradable synthetic polymer. PHAs are produced by a

wide variety of microorganisms as an internal carbon and energy storage, as part of their

27

survival mechanism (Doi, 2002; Jung, 2005). In fact, in vivo studies revealed that PHB

and PHBV, which are members of PHAs family, degrade into d-3-hydroxybutyrate acid,

which is a normal constituent of human blood, which my justify the low toxicity of PHB

(Tokiwa, 2004; Choi, 2005; Cheng, 2006). PHBV is thermoplastic optically active polyester

which is insoluble in water and exhibit a high degree of polymerization (Scandola, 1997;

Kang, 2001). This polymer presents some interesting characteristics, namely antioxidant

(i.e. inhibits the oxidation of other molecules) and piezoelectric properties, i.e. the capacity

of a material to suffer electric polarization due to mechanical stress (Chen, 2000).

Moreover, PHA have various chemical compositions and functionalized groups in the side

chain that allows further chemical modification, which is a major advantage in DDS

development, since it allows to tailor its degradation according with the pretended

application (Wu, 2009; Chen, 2013).

There are several techniques regarding the development of microparticulate systems,

including W/O/W double emulsion, organic phase separation, supercritical fluid, and spray

drying techniques. Among them, the W/O/W double emulsion technique is a well-used

process, which offers the possibility of incorporating both hydrophilic and hydrophobic

molecules, allowing the incorporation of several agents (Freiberg, 2004).

One of aim of the current study concerned the development of microparticles (MPs) made

up of PHBV, a PHB copolymer which presents less crystalline and more flexible, and for

being used as carriers of dexamethasone (Dex) and bovine serum albumin (BSA),

hydrophobic and hydrophilic model drugs, respectively. For this purpose, it was adopted a

water-in-oil-in-water double emulsion solvent evaporation technique, since as previously

mentioned allow the incorporation of water-soluble molecules. The propose method was

previously modified aiming to tailor the delivery of the bioactive agents aiming to

regenerate a defined tissue accordingly with its needs. In this sense, for the production of

PHBV MPs a mixture of different solvents was used in different proportions, namely 100:0

% (v/v) chloroform:ethanol and 90:10%(v/v) chloroform:ethanol. This strategy was

proposed by Poletto and co-workers aiming the control over PHBV particles size using an

emulsification-diffusion technique, and subsequently effect on the bioactive substance

release profile (Poletto, 2008).

When MPs are injected in target tissues, they are continuously exposed to external

mechanical deformation, leading to the uncontrollable displacements of particles

(Fitzgerald, 1987; Griffith, 2000; Chan, 2005). This displacement greatly affects the

concentration of bioactive macromolecules in target tissues, resulting in limited tissue

regeneration (Lemperle, 2004). Moreover, when in free movement the particulate system

28

can be expelled even before release of the therapeutic agent. Therefore, extensive efforts

are being made to resolve this challenge while harnessing the advantages of particulate

systems, including their minimally invasive and controllable drug delivery.

Herein, we propose an innovative delivery system composed of MPs embedded into a

hydrogel that enables the control over release kinetics from simultaneous, combined and

sequential delivery of different biochemical cues in a predefined spatiotemporal manner.

In this sense, the microparticulate system loading bioactive molecules would be combined

with injectable Gellan Gum (GG) hydrogels. Injectable GG hydrogels have been shown to

be a suitable platform to support and deliver cells for noninvasive injectable TE

applications (Oliveira, 2010). With this in mind, a hybrid structure was produce aiming the

co-localization of MPs spatial distribution and consequently the release site, as well as,

producing a long term release that can expedite tissue regenerative processes. In this

sense, BSA and Dex were used as bioactive molecules models; in this way we herein

describe a new approach that allows not only the incorporation and further release of both

relevant bioactive hydrophilic and hydrophobic molecules, as well as their incorporation in

an injectable platform.

29

2. Materials and Methods

2.1. Preparation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles

The PHBV (molecular weight (Mw) = 425,692 g mol-1) (PHB Industrial S.A., Brazil) MPs

were prepared by using a modification of the double emulsion solvent evaporation method

previously described by Ogawa (Ogawa, 1988). Briefly, 3 ml of a 0.5% w/v Poly(vinyl

alcohol) (PVA) (Sigma-Aldrich, Germany) aqueous solution (W1) was mixed with 15 mL of

a 3% w/v PHBV solution in chloroform-ethanol (O). Two different ratios of chloroform-

ethanol were used in order to verify its effect over the release profile of the bioactive

substances, namely 100:0 chloroform:ethanol (100 PHBV MPs) and 90:10

chloroform:ethanol (90 PHBV MPs).The mixture was emulsified for 3 minutes using an

ultra-turrax T18 (IKA-Werke, Germany) in order to obtain the first emulsion (W1/O). The

primary emulsion was then dripped into 150 mL of 1% (w/v) PVA solution (W2) and

homogenized (RW 16 basic IKA-Werke, Germany) during 5 minutes forming the

secondary emulsion (W1/O/W2). The final W1/O/W2 emulsion was left to evaporate under

magnetic stirring for 4–5 h. The obtained microspheres were collected by centrifugation,

washed at least three times with deionized water and freeze-dried.

2.1.1. Bovine Serum Albumin-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

BSA-loaded PHBV MPs were also prepared following the same protocol previously

described. In this sense, BSA (Sigma-Aldrich, Germany) was used at a concentration

0.1% (w/v), being dissolved within the first aqueous phase (W1).

2.1.2. Dexamethasone-loaded Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

A similar procedure described above was followed aiming at incorporate Dex (Sigma-

Aldrich, Germany) within the PHBV MPs. Regarding this, Dex was dissolved in the

organic phase at a final concentration of 3.3x10-3 % (w/v).

2.2. Gellan Gum hydrogel preparation

For the preparation of the injectable GG hydrogels the procedure previously described by

Silva (Silva, 2013) was followed. Gelzan CM (Sigma-Aldrich, Germany) powder was

mixed at room temperature with deionized water at a concentration of 1.25% (w/v) under

30

constant stirring. The solution was heated at 90ºC and kept at this temperature for 30 min.

Subsequently, CaCl2 (VWR, USA) was added to the GG solution at a concentration of

0.18% (w/v) to act as a crosslinker. Finally, the solution was left at room temperature

during approximately one hour, allowing the formation of the gel, which was maintained in

a phosphate-buffered saline (PBS) solution.

2.2.1. Incorporation of Poly (hydroxybutyrate-co-hydroxyvalerate) microparticles in

Gellan Gum hydrogel

To incorporate PHBV MPs into the GG hydrogel, a similar procedure previously described

was followed. In this regard, 30 mg of PHBV MPs were dispersed into 20 mL of a 1.25%

(w/v) GG solution. This solution was then cross-linked with 0.18% (w/v) of CaCl2 and left

at room temperature during approximately one hour, being maintained in a PBS solution.

2.3. Determination of Particles Production Yield

The yield of the particles prepared under the different conditions defined in section 3 was

calculated using Equation 1.

( ) (Equation 1)

The yield was calculated based on the weight of MPs ( ) and compared to the weight of

the compounds used to prepare the MPs ( ), polymer or the sum of polymer and

bioactive substance.

2.4. Efficiency of incorporation of bioactive molecules into the Poly

(hydroxybutyrate-co-hydroxyvalerate) microparticles

2.4.1. Bovine Serum Albumin Quantification

Aiming at determining BSA IE into the MPs, 10 mg of BSA-loaded MPs were completely

dissolved in chloroform with vigorous shaking and at room temperature for 24 hours. This

allows the dissolution of PHBV and at the same time permits the release of the entrapped

BSA. Then 10 mL of PBS were added in order to dissolve the BSA. After centrifugation,

the supernatant was collected and filtered through 0.2 mm membrane. Then, the

concentration of incorporated protein was measured using a Micro BCATM Protein Assay

31

Kit (Pierce, IL) and measured at 562 nm by Synergy HT multi-detection microplate reader

(Bio-Tek's Gen5™, USA). The IE of BSA within PHBV MPs was calculated by equation 2,

as follows:

( )

(Equation 2)

2.4.2. Dexamethasone Quantification

PHBV MPs and incorporated Dex were dissolved as described in section 1. Dex

concentration in the solution was measured at 242 nm by Synergy HT multi-detection

microplate reader in a 96-well quartz plate (n = 3). The incorporation efficiency (IE) of Dex

was calculated as follows:

( )

(Equation 3)

2.5. In vitro Release Studies

2.5.1. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles

In vitro release from PHBV microspheres of both model bioactive substances was

evaluated up to 21 days. Approximately, 30 mg of microspheres were incubated in 10 mL

of PBS (0.01 M, pH = 7.4), and maintained in a precision water bath (Grant, UK) at 60 rpm

and 37±0.5ºC. At defined time points, 1 mL of supernatant was collected and an

equivalent amount of fresh PBS at 37ºC was added to maintain the total volume of the

sample. The concentrations of BSA and Dex in medium were measured at 562 and 242

nm, respectively by using the microplate reader.

2.5.2. In vitro Release Studies from Poly (hydroxybutyrate-co-hydroxyvalerate)

microparticles embedded within Gellan Gum hydrogels

The in vitro release of both bioactive model substances from PHBV MPs embedded within

GG hydrogels was performed during 21 days. Approximately, 220 µL of GG hydrogel

32

embedding PHBV MPs were incubated with 10 mL of PBS (0.01 M, pH = 7.4), and

maintained in a precision water bath at 60 rpm and 37±0.5ºC. At defined time points, 1 mL

of supernatant was collected and an equivalent amount of fresh PBS at 37ºC was added

to maintain the total volume of the sample. The concentrations of BSA and Dex in medium

were measured at 562 and 242 nm, respectively by using the microplate reader.

2.6. Laser Diffraction Spectrometry

The particle size distribution of the prepared MPs was measured by laser diffraction

spectrometry (Coulter LS 230, Coulter Electronics, USA). The dried samples were

suspended in a 0.2% Tween 80 (Sigma-Aldrich, Germany) solution and sonicated for 5

min with an ultra-sound probe (Sonoswiss SW 6 H, Sonoswiss® AG, Switzerland) before

measurement. The obtained homogeneous suspension was used for the particle size

distribution analysis.

2.7. Fourier Transform Infrared spectroscopy

Fourier Transform InfraRed Spectroscopy (FTIR) analyses, under Transmission mode,

were performed in PHBV MPs (unloaded, loaded and after release) in order to analyse

their chemical composition. With this purpose, 1 mg of MPs were mixed with 40 mg of

Potassium Bromide (KBr) and then processed into a disc in a manual press (161-1100

hand press, Pike technologies, Madison, WI). FTIR-KBr spectra (IR Prestige 21,

Shimadzu, Japan) were recorded at 40 scans with a resolution of 4 cm-1, at wavelengths

from 1400 to 4000 cm-1.

2.8. Scanning Electron Microscopy

The morphology of the PHBV MPs and dehydrated PHBV MPs embedded within GG

hydrogels was analysed before and after the in vitro release studies, by scanning electron

microscopy using a model S360 microscope (Leica Cambridge, UK). Moreover, the cross

sections of the MPs were performed after the embedding of this particulate system within

the GG hydrogel which were submitted to several cuts. Before being analysed, the

different structures were gold sputter coated (model SC502; Fisons instruments, UK) for 2

33

minutes at 15 mA. Micrographs were recorded at 5.0 kV with magnifications of 250, 1 000,

2 000, 8 000 and 20 000.

3. Results

3.1. Poly (hydroxybutyrate–co–hydroxyvalerate) microparticles Process Production

The PHBV MPs were produced following a double emulsion solvent evaporation method

(Figure 1). The propose method was previously modified with the objective of tailoring the

delivery of relevant biochemical cues aiming to regenerate a defined tissue accordingly

with its needs. In this sense, for the production of PHBV MPs a mixture of different

solvents was used in different proportions, namely 100:0 (v/v) chloroform:ethanol and

90:10 (v/v) chloroform:ethanol.

Figure 1 – Schematic depiction of double emulsification evaporation method.

After the preparation of PHBV MPs, the process yield was calculated for the particles

obtained under different conditions aiming to assess its reproducibility, as well as, the

potential effect of the tested conditions (Table I). The results showed that when ethanol

was added to the process, yield decreased 6%.

Table I. Yield and Size of the PHBV MPs produced under different conditions.

Processing Conditions

Chloroform:Ethanol ratio

Yield, %

(mean ±SD)

Particle Size

Distribution, µm

(mean ±SD) PHBV (100:0) 95.7 ±2.32 41.9 ±20.2

PHBV (90:10) 90.3 ±2.43 57.6 ±29.1

34

The size distributions of the MPs obtained using different proportions of

chloroform/ethanol in the organic phase for the first emulsification was further determined

and the results are shown in Table I. PHBV MPs prepared exclusively with chloroform

showed a mean particle size range of 15.6 to 76.4 µm, whilst the particles obtained using

ethanol present a particle size range of 5.6 to 91.1 (Figure 2).

Figure 2 – Particle size distribution of the PHBV MPs.

The mean particle size of unloaded-MPs obtained by adding ethanol to the process was

found to be larger than MPs where it was exclusively used chloroform.

The morphology of the produced MPs was analysed by SEM (Figure 3). Perfect spherical

shape MPs of different sizes were observed (Figure 3 A1 and B1) which is in accordance

with the measured particle size distribution (Table I). Moreover, a detailed analysis of MPs

surface topography revealed rough surfaces with the presence of some micro-size pores

(Figure 3 A3 and B3). The presence of this porosity is not dependent on the microsphere

preparation conditions, although, the addition of ethanol leads to a different surface

topography.

35

Figure 3 - SEM micrographs of PHBV MPs obtained under different experimental conditions: (A) using 100% chloroform; and (B) a mixture of 90% chloroform: 10% ethanol. The sequential micrographs represent an overview of the obtained MPs showing the detail of their surface and respective close up of the internal morphology.

50 µm 50 µm

B1

10 µm

A2

10 µm

B2

A3

5 µm

B3

5 µm

10 µm

A4

10 µm

B4

A1

36

Like for the external porosity, independently of the processing conditions used, the

produced MPs presented a porous core (Figure 3A4 and B4). Nonetheless, the addition of

ethanol to the organic phase leads to the formation of larger pore (Figure 3B4).

3.2. Bovine Serum Albumin and Dexamethasone Entrapment Efficiency and Release

Profile

3.2.1. Loaded Microparticles properties

The impact of the incorporation of bioactive substances within PHBV MPs on the process

yield and particle size was further determined (Table II).

Table II. Yield and Size of the PHBV loaded MPs produced under different conditions.

Processing Conditions

Chloroform:Ethanol ratio

Yield, %

(mean ±SD)

Particle Size Distribution, µm

(mean ±SD)

Entrapped bioactive

molecule Dex BSA Dex BSA

PHBV (100:0) 95.4 ±1.36 88.8 ±5.50 55.0 ±39.4 78.5 ±78.5

PHBV (90:10) 89.9 ±5.64 79.0 ±9.83 58.3 ±27.5 63.0 ±42.2

Regarding the process yield results, no significant differences were observed between the

unloaded-PHBV MPs (Table I) and Dex-loaded MPs (Table II). However, the results also

showed that BSA-loaded MPs presented a lower yield than the unloaded-MPs.

In what concerns the particle size of the PHBV-loaded MPs, the incorporation of the

bioactive molecules lead to an increased size, with the exception of Dex in the presence

of ethanol (Table II). Furthermore, increase in size of MPs was found to be proportional to

the molecular weight of the bioactive substance that was incorporated. Thus, the MPs

loaded with Dex (392.46 Da) exhibited a reduced mean particle size when compared with

the MPs loaded with BSA (approx. 66 kDa) (Table II).

The SEM analysis of the loaded-PHBV MPs confirmed not only the particles size

measurements but also the maintenance of the spherical morphology after incorporation

of the bioactive molecules (Figure 4A1 and B1). Despite, the morphology of the loaded

MPs surface was similar to the surface of the unloaded-PHBV MPs produced in the

37

absence of ethanol (Figure 3A3). These observations were independent of the

incorporated molecule (Figure 4A2 and B2).

Figure 4 - SEM micrographs of BSA loaded-PHBV MPs obtained under different experimental

conditions: (A) using 100% chloroform; and (B) 90% chloroform: 10% ethanol. The sequential

micrographs represent an overview of the obtained MPs showing the detail of their surface.

The incorporation efficiency of BSA and Dex within the PHBV MPs obtained under

different experimental conditions are described in Table III.

50 µm

A1

50 µm

B1

10 µm

A2

10 µm

B2

A3

5 µm

B3

5 µm

38

Table III. Incorporation Efficiency of Bovine Serum Albumin and Dexamethasone within PHBV

MPs.

Processing Conditions Chloroform:Ethanol ratio

Incorporation Efficiency, %

(mean ±SD)

Dex BSA

PHBV (100:0) 81.6 ±8.41 66.2 ±3.33

PHBV (90:10) 92.3 ±2.17 65.5 ±1.72

The entrapment efficiency results of BSA inside PHBV MPs revealed no significant

differences when ethanol was added to the organic phase and efficiency around 66% was

achieved independently of the processing conditions (Table III).

Regarding the incorporation efficiency of Dex, it was found to be 81.6 ±8.41% in the

absence of ethanol on the organic phase, while the other was 92.3 ±2.17%. Thus, the

bioactive molecule entrapment efficiency was affected by organic phase composition.

The entrapment of BSA within PHBV MPs was further confirmed by FTIR analysis (Figure

5). The FTIR spectra of BSA-loaded PHBV MPs revealed the BSA’ characteristic bands

for both formulations, indicating that the hydrophilic drug was successfully incorporated

within the PHBV MPs.

Figure 5 – FTIR spectra of BSA and PHBV MPs: BSA (black line); 100 unloaded-PHBV MPs (red

line); 90 unloaded-PHBV MPs (blue line); 100 BSA-loaded PHBV MPs (pink line); 90 BSA-loaded

PHBV MPs (green line). The characteristics bands of BSA are marked (*).

39

FTIR analysis was also performed aiming to confirm the Dex entrapment inside PHBV

MPs (Figure 6). The FTIR spectra of Dex-loaded PHBV MPs showed its characteristics

bands, indicating the effectiveness of the entrapment procedure of Dex inside the MPs

obtained by using different organic phase compositions.

Figure 6 – FTIR spectra of Dex and PHBV MPs: Dex (black line); 100 unloaded-PHBV MPs (red

line); 90 unloaded-PHBV MPs (blue line); 100 Dex-loaded PHBV MPs (pink line); 90 Dex-loaded

PHBV MPs (green line). The characteristics bands of Dex are marked (*).

3.2.2. Bovine Serum Albumin Release from Poly (hydroxybutyrate–co–

hydroxyvalerate) microparticles

The release profile of BSA from BSA-loaded PHBV MPs obtained under different

conditions and up to 21 days is shown in Figure 7. The results of the MPs produced in the

absence of ethanol indicate that 76% of BSA was released within the first 8 hours of

incubation reaching a release of approximately 78% of the incorporated BSA at day 21. In

contrast, the BSA loaded-MPs produced with a mixture of chloroform and ethanol, as

organic phase, released about 78% of the incorporated BSA within the first 8 hours. After

21 days of incubation, approximately 80% of BSA was released from BSA-loaded PHBV

MPs.

The BSA release profiles were confirmed by FTIR analysis (Figure 8). After the release

studies, the FTIR spectra of BSA-loaded 100 PHBV MPs, in which the organic phase

40

does not contain ethanol, revealed a reduction in the intensity of the BSA characteristic

bands (Figure 8 – dark blue line). However, it is evident the presence of the typical peaks

of BSA even after 21 days of incubation from BSA-loaded 100 PHBV MPs.

Figure 7 - In vitro release profile of BSA from BSA-loaded PHBV MPs obtained using a mixture of

chloroform and ethanol in different proportions: (100:0) % ( ); and (90:10) % ( ).

Figure 8 – FTIR spectra of BSA-loaded PHBV MPs: BSA-loaded 100 PHBV MPs (black line) and

BSA-loaded 90 PHBV MPs (red line) before release studies; and BSA-loaded 100 MPs (blue line)

and BSA-loaded 90 MPs (purple line) after 21 days of in vitro release. The characteristics bands of

BSA are marked (*).

41

The degradation of the MPs incorporating both hydrophobic and hydrophilic bioactive

substances (BSA and Dex) was monitored by SEM. As shown in Figure 9, there are

slightly morphological and topographical differences after the in vitro release studies.

However, is possible to observe for almost all the formulations that the microparticulate

systems maintain their structure and integrity.

Figure 9 - SEM micrographs of BSA-loaded PHBV MPs obtained under different experimental conditions: (A) using 100 % chloroform; and (B) with 90% chloroform: 10% ethanol, after 21 days of in vitro release studies. The sequential micrographs represent an overview of the obtained MPs showing the detail of their surface.

3.2.3. Dexamethasone Release from Poly (hydroxybutyrate–co–hydroxyvalerate)

microparticles

The release of Dex from Dex-loaded PHBV MPs obtained under different conditions and

up to 21 days is shown in Figure 10. The results from the Dex loaded-100 PHBV MPs

produced in the absence of ethanol indicate that 55% of Dex was released within the first

8 hours of incubation reaching a release of approximately 69% of the incorporated Dex at

day 21. In opposition, the Dex loaded-MPs produced with a mixture of chloroform and

ethanol as organic phase released about 89% of the incorporated Dex within the first 8

hours and no Dex was left in the MPs after 5 days of incubation.

50 µm

A1

10 µm

A2

5 µm

A3

50 µm

B1

10 µm

B2

5 µm

B3

42

The Dex release profiles were confirmed by FTIR analysis (Figure 11). The FTIR spectra

of Dex-loaded 90 PHBV MPs after the release studies revealed a reduction in the intensity

of the Dex characteristic bands in the IR spectrum of the Dex-loaded PHBV MPs wherein

the organic phase consists only of chloroform (Figure 11 – dark blue line) and the

absence of those bands for the Dex loaded-100 PHBV MPs.

Figure 10 - In vitro release profile of Dex loaded-MPs obtained using a mixture of chloroform and ethanol in different proportions: (100:0) % ( ); and (90:10) % ( ).

Figure 11 – FTIR spectra of Dex-loaded PHBV MPs: Dex-loaded 100 PHBV MPs (black line) and Dex-loaded 90 PHBV MPs (red line) before release studies; and 100 Dex-loaded MPs (line) and 90 Dex-loaded MPs (purple line) after 21 days of in vitro release. The characteristics bands of Dex are marked (*).

43

3.3. Injectable Gellan Gum/ Poly (hydroxybutyrate–co–hydroxyvalerate)

microparticles System

3.3.1. Injectable System Properties

The main goal of this work is to propose a stable DDS based on the obtained PHBV MPs.

For this purpose, the different PHBV MPs were embedded within the GG’ solution prior

gelation.

Figure 12 – Schematic representation of injectable fluorescein isothiocyanate labelled bovine serum albumin – loaded PHBV MPs embedded within GG hydrogel system.

The SEM analysis of the system after gelation confirmed the homogeneous dispersion of

the MPs with the hydrogel without losing their spherical shape (Figure 13).

Figure 13 - SEM micrographs of PHBV MPs embedded within injectable GG hydrogels. The sequential images represent an overview of the distribution of the particles within the hydrogel, close up on the surface of the biphasic structure, and a cross-section.

400 µm

A1

100 µm

A2

400 µm

A3

44

Moreover, PHBV-MPs were completely embedded with the GG matrix guaranteeing its

delivery as a stable combined system.

3.3.2. Bovine Serum Albumin Release from the Injectable Gellan Gum/ Poly

(hydroxybutyrate–co–hydroxyvalerate) microparticles System

In order to assess the impact of the incorporation of the MPs within the hydrogel over BSA

release profile, release studies under the conditions defined for the single MPs were

carried out (Figure 14).

Figure 14 - In vitro release profile of BSA from GG hydrogels embedding PHBV MPs obtained using a mixture of chloroform and ethanol in different proportions: (100:0) % ( ); and (90:10) % (

).

The release profile confirmed that the presence of the GG hydrogel surrounding the MPs

reduced significantly the initial burst effect. The results of the MPs produced in the

absence of ethanol embedded within GG hydrogel indicate that 35% of BSA was released

within the first 8 hours of incubation reaching a release of approximately 36% of the

incorporated BSA at day 21. In contrast, the BSA loaded-MPs produced with a mixture of

chloroform and ethanol as organic phase combined with the GG matrix released about

40% of the incorporated BSA within the first 8 hours and after 21 days of incubation

approximately 41% of BSA was released from this system.

After 21 days of immersion in PBS the system remained stable in the sense that PHBV

MPs were still embedded in the system (Figure 15). Nonetheless, the hydrogel revealed

45

some signs of degradation, the surface erosion was apparent causing the formation of

some cavities and suggesting the loss of particles into the medium (Figure 15A2).

Figure 15 - SEM photographs of BSA loaded-MPs embedded within GG hydrogel, after 21 days of in vitro release studies. The sequential images represent an overview of the distribution of the particles within the hydrogel, close up on the surface of the biphasic structure, and a cross-section.

3.3.3. Dexamethasone Release from the Injectable Gellan Gum/Poly

(hydroxybutyrate–co–hydroxyvalerate) microparticles System

Regarding the Dex-loaded PHBV MPs embedded within GG hydrogels, the effect of the

GG matrix on the release profile was evaluated under release studies up to 21 days

(Figure 16).

Figure 16 - In vitro release of Dex from GG hydrogels embedding PHBV MPs obtained using a

mixture of chloroform and ethanol in different proportions: (100:0) % ( ); and (90:10) % ( ).

400 µm

A1

100 µm

A2

400 µm

A3

46

The release profile confirmed that GG hydrogel surrounding the MPs eliminated the initial

burst. Dex release from MPs entrapped into the GG hydrogel showed a nearly zero-order

release profile. Moreover, Dex was not dependent on the MPs processing conditions as

no differences were observed between the systems produced in the presence and in the

absence of ethanol. Both systems attained a release of around 26% at the end of 21 days

of incubation in PBS.

4. Discussion

From the clinical point of view, injectable hydrogels formed by in situ polymerization are

highly desirable since they gain advantages over preformed hydrogels: (1) enabling

minimally invasive surgeries for implantation; (2) formation in any desired shape in good

alignment with surrounding tissue defects; and (3) easy encapsulation of bioactive

molecules and cells (Jin, 2012). With a view toward developing a minimally-invasive DDS

that would provide sustained local release of biochemical cues, we designed a biphasic

injectable DDS to be applicable as a long term release system with two distinct objectives:

i) the control over release profile of proteins and/or glucocorticoids and ii) the spatial

distribution of the particles avoiding its uncontrolled migration through the human body.

For this purpose, BSA- and Dex-loaded PHBV MPs, processed by a double

emulsification-solvent evaporation method with modifications, were incorporated into

injectable GG hydrogels. In the future, this system will allow overcoming the

uncontrollable displacements, as well as, the clearance of particles by the organism.

In this study, PHBV a biodegradable thermoplastic polymer was chosen due to its singular

properties, namely biocompatibility, low toxicity, and slow degradation. In his turn, GG was

used because of its stability, biocompatibility and biodegradability. BSA and Dex were

respectively used as hydrophilic/proteins and hydrophobic/glucocorticoids model

molecules envisioning the possibility of a combined localized administration of an anti-

inflammatory agent and a cytokine or growth factor by controlled release where high local

concentrations are needed and/or local concentration gradients need to be established for

successful tissue regeneration (Chen, 2010).

While developing microparticulate DDS there are several aspects that are critical since it

is a complex procedure that involves several processing and design variables. Even

slightly changes of variables and system components can have significant impact on

unique features of MPs including particle size and morphology, as well as entrapment

efficiency of the molecules of interest. The double emulsion solvent evaporation method

47

described by Ogawa et al. (1988), is a highly reproducible method that has been

extensively used to create particles with a spherical structure and inner core-shell

structure. MPs of PHBV have been previously produced by this methodology (Li, 2009).

However, in order to optimize the reported methodology aiming to tailor the bioactive

molecules release profile, a mixture of different solvents, as organic phase, is herein

proposed. Chloroform was selected as the solvent for the PHBV and ethanol to improve

the mutual solubility of the organic and the aqueous phases and based on the type of

molecular interactions between the organic solvents (dipole-dipole) and between ethanol

and water (hydrogen bond) (Poletto, 2008).

The appropriate size of MPs is an important parameter that affects the release behaviour

of the bioactive substances. The produced PHBV MPs showed a mean particle size

varying from 41.9 to 78.5 µm, which was affected by the processing conditions, presence

of ethanol, as well as by the type of bioactive molecule incorporated. The size of the MPs

in the presence of ethanol was higher than when only chloroform was used as organic

phase. These results are not in accordance to what was previously reported by Poletto

(2008) that showed a decrease in the size of PHBV nanoparticles when ethanol was used

as a surface agent. The distinct results are probably due to the differences between both

methods used to produce the particles; while they followed an emulsification-diffusion

technique, the double emulsion solvent evaporation method herein proposed involves

intense high–speed homogenization, proved to have a tremendous effect on the particle

size (Mukherjee, 2008). Additionally, it is important to note that the diameters of these

MPs are less than 100 µm, and thus can be readily injected into the body with comparable

less pain (Mathew, 2007; Wang, 2007).

The effect of employing a binary mixture of solvents as organic phase was also

investigated in terms of morphology. The MPs showed a well-defined spherical shape

without visual defects, independently of the processing conditions showing a considered

particle size variation, which is in accordance with the particle size distribution results

already reported. When employed emulsion/solvent evaporation method, the roughness

pattern on the microspheres surface has been associated to the high crystallinity and fast

precipitation of this polymer, after solvent evaporation from the internal phase of the

emulsion (Bidone, 2009). However, the comparison of the results revealed that the

surface topographies were also influenced by the addition of ethanol to the process.

Moreover, the presence of micro-size pores on PHBV microspheres surfaces,

independently of the microsphere preparation conditions, was observed. The pores could

help to accelerate the release of the bioactive molecules from the microspheres.

Additionally, both particulate system obtained under different experimental conditions

48

possess highly porous internal structure, however the addition of ethanol on the organic

phase lead to larger core pores. This could be attributed to the hydrophobicity of PHBV;

during the first emulsification process, water droplets (first aqueous phase) can interact

with the ethanol present in the organic phase, the same is not observed when the organic

phase is composed only by chloroform, leading to the formation of larger droplets inside

the dispersed phase and consequently to larger the pores after removal of the water

droplets by freeze-drying (Zhu, 2007).

In addition to the processing conditions, the incorporation of BSA and Dex within the

PHBV MPs also affected the properties of the produced MPs. An increase in mean size

was observed for the PHBV loaded MPs, which was proportional to the molecular weight

of the incorporated bioactive molecule. The BSA-loaded MPs (BSA has a MW of 66 kDa)

exhibited a larger mean size in comparison to Dex-loaded MPs (Dex has a MW of 392.46

Da). Studies about the inner distribution of hydrophilic and hydrophobic dyes-loaded

polymeric MPs, synthesized by double emulsion method and characterized by confocal

microscopy, have showed a localization of hydrophilic dye in the core of particles and the

hydrophobic dyes in the shell structure of MPs (Lamprecht, 2000; Mao, 2007).

The entrapment efficiency of Dex inside PHBV MPs was found to be influenced by the

presence of ethanol in the organic phase during particle preparation. Based on the results,

the incorporation efficiency of Dex was higher when adding ethanol to the process what

can be explained by the dipole-dipole interaction with chloroform and its hydrogen bond

with the water. Moreover, we hypothesize that the presence of the ethanol could promote

the entrapment of Dex in the core, in similarity with what happens with the hydrophilic

substances, by binding its apolar part to Dex while its hydrogen bond was connected to

the water. Regarding the protein, BSA incorporation efficiency revealed to be independent

of the organic phase composition. In this sense, comparing the results of both model

bioactive molecules it is possible to observe that Dex has much higher incorporation

efficiency than protein. As commonly seen in the double emulsion method, the

incorporation of bioactive molecules is highly dependent on the chemical structure and

molecular weight of the molecules, and hydrophobicity of the polymers (Park, 2005).

Moreover, the PHBV hydrophobicity may not be compatible with the hydrophilicity of the

protein, what can lead to a formation of an instable emulsion. Thus, the bioactive molecule

entrapment efficiency was affected by organic phase composition and the incorporated

active substance molecular weight. Lionzo et al. proposed a general single emulsion

solvent evaporation technique for the incorporation of Dex within PHBV microspheres and

the results appeared to be very similar with the results herein reported (Lionzo, 2007).

49

The in vitro release profile from a polymeric matrix is controlled by a variety of factors,

namely the solubility of the drug within the surrounding media, the molecular weight of the

bioactive substance, its mobility within the swollen polymeric network, as well as the

dissolution rate of the polymer and polymer-drug interactions (Balmayor, 2009).

Nevertheless, there are some other features that can have influence in the release kinetic,

namely morphological characteristics of the particulate carriers, physicochemical

characteristics of the polymer, and particle size and size distribution, which can be

controlled, as previously described, by varying the conditions of the particles preparation

procedure (Embleton, 1993; Saha, 1994; Igartua, 1997; Yang, 2001).

Although PHAs have been described as high crystalline polymers (the reason for its slow

degradation), excessive release rates from PHAs microspheres have been previously and

consistently reported and assumed that the release phenomenon is more dependent on

drug dissolution rather than matrix degradation or diffusion (Bidone, 2009). This was

confirmed by the microscopic analysis of the MPs after the release studies that showed

alterations in the surface of the systems but structurally intact and with spherical shape.

Moreover, Independently of the conditions and of the incorporated molecules, the release

profiles showed an initial burst phase, followed by a slower release, typical of a first order

release kinetics. This behaviour is in accordance to what was described for other

biodegradable polymers (Tuncel, 1995; Zignani, 1997; Hanes, 1998; Huang, 2001;

Vandamme, 2002; Yoon, 2003). The in vitro release results indicated that more than 68%

of Dex was released after 21 days from the particles produced in the absence of ethanol

and 100% of Dex after 5 days from the particulate system produced in the presence of

this solvent. Regarding the BSA release studies, no significant differences were observed

when ethanol was added to the organic phase, and a release of protein around 78% was

observed after 21 days of immersion. Accordingly with these observations, in the case of

MPs produced with the absence of ethanol in the organic phase, the bioactive molecules

release profile present a more sustained pattern when compared with the MPs where it

was added ethanol. This finding, was expected since the addition of this alcohol to the

organic phase creates larger pore size what can explain the higher drug release profile,

since their presence facilitate the water penetration.

Envisioning an administration improved by more precise spatial and temporal placement

of the MPs within the organism, we propose the combination of the previous system with

an injectable hydrogel. Injectable GG hydrogels have been extensively explored within our

Research Group for a variety of TE and RM applications, by taking advantage of their

thermal and ionic sensitivity thus allowing its gelation in situ (Oliveira, 2010; Coutinho,

2010; Pereira, 2011).

50

In this sense, the incorporation of BSA- and Dex- loaded PHBV MPs within GG hydrogels

envisioned as a long term release system with predefined control over release of proteins

and/or glucocorticoids was developed. This biphasic system showed a uniform distribution

of PHBV MPs across the GG matrix with a strong connection with the structure. However,

the incorporation of the MPs into the GG hydrogels strongly influenced the profile and the

released amount of both Dex and BSA. The previously observed differences with the

processing conditions were reversed and the initial burst effect was eliminated. The

microspheres entrapped into GG hydrogels showed a nearly zero-order release profile,

which indicated that the combination of both systems causes a sustained release pattern.

This phenomenon can be attributed to the high density of GG, which must limit the

bioactive molecules diffusion. Additionally, the disintegration rate of the GG hydrogels was

also verified after immersed in PBS for 3 weeks. It was observed that the hydrogel

revealed some signs of degradation on the surface, which is in accordance with what was

reported by Chang (2010). This evidence can be explained by hydrolytic reactions, being

possibly exacerbate by the PHBV microspheres acidic degradation products.

In summary, this work proved the capacity of PHBV microparticulate systems as a unique

tool to deliver both hydrophilic and hydrophobic molecules, and the effect of the organic

phase composition on their incorporation and respective release profile. Moreover, the

embedding of this system with injectable hydrogel structures reinforced that the proposed

system constitutes a suitable platform for being used as a sustained release system

allowing, at the same time the modulation of the localization of particles obtained. The

biphasic DDS herein described demonstrate the ability to deliver multiple bioactive

molecules with distinct kinetics, in real-time, which is likely required to drive tissue

development to regeneration. Furthermore, these systems also minimize the release of

bioactive molecules, which constitutes an advantage considering that tissue regeneration

normally occurs over long period of time frames (Biondi, 2008).

5. Conclusion

The main goal of this work was to develop a simple and versatile delivery platform, based

on the combination of microparticulate systems and an injectable hydrogel, to carry

hydrophilic and hydrophobic model bioactive substances relevant for TE purposes. In this

sense, an injectable Gellan Gum hydrogel entrapping bioactive molecules-loaded PHBV

microspheres has been prepared and characterized. Such platform showed the ability for

being used as a long term release structure, which constitutes an advantage considering

that tissue regeneration normally occurs over long period of time frames, offering at the

51

same time a high spatial control over polymeric particles distribution at the injection site.

The proposed system provides a promising strategy for a variety of localized controlled

drug delivery applications, namely in soft TE applications.

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IV. Final Remarks and Future Perspectives

The main goal of this work was to develop a simple and versatile delivery platform, based

on the combination of microparticulate systems, carrying model bioactive substances

relevant for Tissue Engineering and Regenerative Medicine purposes, and an injectable

hydrogel. Such platform was expected to comprise features that would allow its

application as a long term release DDS, offering a high spatial control over and enhanced

resident time of polymeric particles upon injection.

PHBV MPs were produced by means of a double emulsion solvent evaporation technique

based on a previously reported methodology but with modifications. A mixture of

chloroform and ethanol as organic phase was used with the expectation of tailoring the

release profile of incorporated molecules by changing the physicochemical characteristics

of PHBV microparticles. As expected, the addition of ethanol to the organic phase had

great influence onto PHBV MPs surface topography, size distribution and surface and

core porosity. Moreover, the presence of ethanol in the organic phase benefitted Dex

incorporation but did not affect the entrapment efficiency of BSA. Thus, in addition to the

processing conditions, the type and size of the molecules to be incorporated, in which Dex

and BSA differ significantly, also influence the efficiency of loading. In his turn, the in vitro

release profile studies confirmed an initial burst effect which was followed by a sustained

pattern, typical of a first order release kinetics and in accordance to what has been shown

for microparticulate polymeric DDS.

An innovative approach that tackled the limitations of current microparticulate DDS when

injected as a suspension into the site of injury, namely the reduced residence time and

uncontrolled mechanical forces to which are subjected, was envisioned. The use of

injectable GG hydrogel formulations, have been proposed for diverse tissue engineering

applications and thus would not negatively affected host response, when considered as

MPs carriers. A combined and well integrated system of PHBV MPs within a GG hydrogel

was successfully achieved. This biphasic system was capable of long-term retention, in

comparison to single PHBV microspheres, of the incorporated bioactive substances,

remaining substantially intact.

In this sense, a biphasic DDS, which allows the release of biochemical cues with different

physicochemical features, as well as its localized deliver, was successfully developed and

represents a versatile tool to prepare instructive cell microenvironments for TERM.

60

Further improvements can be foreseen in order to tune the release profile of the combined

structures according to tissue and pathology/injury specific requirements. This will be

attempted by tailoring the properties of the hydrogel, which has been extensively studied

in our Research Group. Moreover, studies to assess systems sensitivity to enzyme-driven

degradation will be designed as it will certainly contribute for modulating the release

kinetics of incorporated molecules.

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V. Annex

1. Morphological Analysis

The SEM analysis of the Dex loaded-PHBV MPs show a similar spherical morphology as

observed in the BSA-loaded PHBV MPs.

Figure 1 - SEM micrographs of Dex loaded-PHBV MPs obtained under different experimental

conditions: (A) using 100% chloroform; and (B) 90% chloroform: 10% ethanol. The sequential

micrographs represent an overview of the obtained MPs showing the detail of their surface.

10 µm

A1 B1

50 µm 50 µm

A2

A3

B2

B3

10 µm

5 µm 5 µm

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63