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Development of Nanocomposite Corticosteroid Particles for use in Asthma

Permanent Link: http://ufdc.ufl.edu/UFE0041619/00001

Material Information

Title: Development of Nanocomposite Corticosteroid Particles for use in Asthma
Physical Description: 1 online resource (117 p.)
Language: english
Creator: Patel, Gina
Publisher: University of Florida
Place of Publication: Gainesville, Fla.
Publication Date: 2010

Subjects

Subjects / Keywords: asthma, chitosan, corticosteroids, dry, glucocorticoids, inhalation, inhaler, nanoparticles, pla, plga, powder, spray
Pharmaceutics -- Dissertations, Academic -- UF
Genre: Pharmaceutical Sciences thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

Notes

Abstract: Asthma is a chronic inflammatory condition of the airways resulting in episodes wheezing and breathlessness. Inhaled Corticosteroids (ICS) are used to prevent the occurrence of asthma attacks. The clinical effect of ICSs depends on the time the drug resides in the lung ? therefore increasing the drug residence time in the lung improves asthma therapy. It has been proposed that nanoparticles could escape clearance mechanisms in the lung and adhere strongly to the lung surface, leading to increased residence time. There are two main barriers to this approach, firstly, nanoparticles cannot deposit in the lung, and instead they are exhaled. Secondly, the particles must be formulated to release drug slowly in order to take advantage of the increased residence time. In order to further improve lung targeting of corticosteroids, poly(lactic-co-glycolic acid) PLGA and poly(lactic acid) PLA were used to produce polymeric nanoparticles of Triamcinolone Acetonide (TA) and Mometasone Furoate (MF) using the solvent evaporation technique. TA was used as a model corticosteroid to allow optimization of various production parameters to produce PLGA nanoparticles. The solvent evaporation technique was used to produce polymeric nanoparticles. A number of process parameters may be varied to alter nanoparticle size and drug encapsulation efficiency. Polyvinyl alcohol (PVA) surfactant concentration and PLGA content were varied to determine their influence on particle size, drug encapsulation efficiency, particle morphology and in vitro release characteristics of TA nanoparticles. Increasing PLGA content resulted in a trend of increasing particle size and drug encapsulation. As PVA concentration was increased particles tended to reduce in size and drug loading. Nanoparticles produced ranged in particle size between 156-209nm. In addition when a low concentration of 1% or 2% w/v PVA was used to produce nanoparticles combined with the use of only 200 mg PLGA, TA crystals were observed by scanning electron microscopy. In vitro release studies revealed TA-PLGA nanoparticles released drug at a similar rate to micronized TA, with 50% drug release being observed within 15 minutes. In order to produce nanoparticles which can deliver drug at a slower rate compared to micron sized particles, a number of changes can be made to nanoparticle production; a more lipophilic drug MF in combination with a more hydrophobic polymer, PLA can be used to further slow down drug release. Subsequently MF nanoparticles (MF-PLA) were using 10 mg MF, 400 mg PLA and 1% w/v PVA, these particles showed slow release compared to MF contained in the Asmanexregistered trademark TwisthalerTM. To further reduce MF release rate, nanoparticles were coated with chitosan. In vitro release studies showed that chitosan coated MF-PLA nanoparticles (CH-MF) showed significantly slower release compared to both uncoated nanoparticles and MF contained within the Asmanexregistered trademark TwisthalerTM. In vitro release studies determined 100% MF occurred after 1 hour for the Asmanexregistered trademark formulation, in comparison at this time only 50% and 24% MF was released from MF-PLA and CH-MF respectively. A novel spray dryer was designed so that operating conditions of this device allowed outlet temperatures to remain below the glass transition temperature of PLGA and PLA, thus making this system suitable to spray dry polymeric nanoparticles. Incorporation of these nanoparticles into lactose based microspheres by spray drying resulted in spherical nanocomposite microspheres. Analysis of these microspheres showed complete incorporation of nanoparticles into the formulation. Optimal conditions for incorporation of nanoparticles into microspheres were using a composition of 75% nanoparticles and 25% lactose. The fine particle fraction of the microspheres was comparable to that of MF from the Asmanexregistered trademark TwisthalerTM. A biphasic release of MF was observed from the microspheres, with a significantly slower release compared to Asmanexregistered trademark. The spray drying process did not seem to alter the release properties of chitosan coated nanoparticles.
General Note: In the series University of Florida Digital Collections.
General Note: Includes vita.
Bibliography: Includes bibliographical references.
Source of Description: Description based on online resource; title from PDF title page.
Source of Description: This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility: by Gina Patel.
Thesis: Thesis (Ph.D.)--University of Florida, 2010.
Local: Adviser: Hochhaus, Guenther.
Electronic Access: RESTRICTED TO UF STUDENTS, STAFF, FACULTY, AND ON-CAMPUS USE UNTIL 2011-08-31

Record Information

Source Institution: UFRGP
Rights Management: Applicable rights reserved.
Classification: lcc - LD1780 2010
System ID: UFE0041619:00001

Permanent Link: http://ufdc.ufl.edu/UFE0041619/00001

Material Information

Title: Development of Nanocomposite Corticosteroid Particles for use in Asthma
Physical Description: 1 online resource (117 p.)
Language: english
Creator: Patel, Gina
Publisher: University of Florida
Place of Publication: Gainesville, Fla.
Publication Date: 2010

Subjects

Subjects / Keywords: asthma, chitosan, corticosteroids, dry, glucocorticoids, inhalation, inhaler, nanoparticles, pla, plga, powder, spray
Pharmaceutics -- Dissertations, Academic -- UF
Genre: Pharmaceutical Sciences thesis, Ph.D.
bibliography   ( marcgt )
theses   ( marcgt )
government publication (state, provincial, terriorial, dependent)   ( marcgt )
born-digital   ( sobekcm )
Electronic Thesis or Dissertation

Notes

Abstract: Asthma is a chronic inflammatory condition of the airways resulting in episodes wheezing and breathlessness. Inhaled Corticosteroids (ICS) are used to prevent the occurrence of asthma attacks. The clinical effect of ICSs depends on the time the drug resides in the lung ? therefore increasing the drug residence time in the lung improves asthma therapy. It has been proposed that nanoparticles could escape clearance mechanisms in the lung and adhere strongly to the lung surface, leading to increased residence time. There are two main barriers to this approach, firstly, nanoparticles cannot deposit in the lung, and instead they are exhaled. Secondly, the particles must be formulated to release drug slowly in order to take advantage of the increased residence time. In order to further improve lung targeting of corticosteroids, poly(lactic-co-glycolic acid) PLGA and poly(lactic acid) PLA were used to produce polymeric nanoparticles of Triamcinolone Acetonide (TA) and Mometasone Furoate (MF) using the solvent evaporation technique. TA was used as a model corticosteroid to allow optimization of various production parameters to produce PLGA nanoparticles. The solvent evaporation technique was used to produce polymeric nanoparticles. A number of process parameters may be varied to alter nanoparticle size and drug encapsulation efficiency. Polyvinyl alcohol (PVA) surfactant concentration and PLGA content were varied to determine their influence on particle size, drug encapsulation efficiency, particle morphology and in vitro release characteristics of TA nanoparticles. Increasing PLGA content resulted in a trend of increasing particle size and drug encapsulation. As PVA concentration was increased particles tended to reduce in size and drug loading. Nanoparticles produced ranged in particle size between 156-209nm. In addition when a low concentration of 1% or 2% w/v PVA was used to produce nanoparticles combined with the use of only 200 mg PLGA, TA crystals were observed by scanning electron microscopy. In vitro release studies revealed TA-PLGA nanoparticles released drug at a similar rate to micronized TA, with 50% drug release being observed within 15 minutes. In order to produce nanoparticles which can deliver drug at a slower rate compared to micron sized particles, a number of changes can be made to nanoparticle production; a more lipophilic drug MF in combination with a more hydrophobic polymer, PLA can be used to further slow down drug release. Subsequently MF nanoparticles (MF-PLA) were using 10 mg MF, 400 mg PLA and 1% w/v PVA, these particles showed slow release compared to MF contained in the Asmanexregistered trademark TwisthalerTM. To further reduce MF release rate, nanoparticles were coated with chitosan. In vitro release studies showed that chitosan coated MF-PLA nanoparticles (CH-MF) showed significantly slower release compared to both uncoated nanoparticles and MF contained within the Asmanexregistered trademark TwisthalerTM. In vitro release studies determined 100% MF occurred after 1 hour for the Asmanexregistered trademark formulation, in comparison at this time only 50% and 24% MF was released from MF-PLA and CH-MF respectively. A novel spray dryer was designed so that operating conditions of this device allowed outlet temperatures to remain below the glass transition temperature of PLGA and PLA, thus making this system suitable to spray dry polymeric nanoparticles. Incorporation of these nanoparticles into lactose based microspheres by spray drying resulted in spherical nanocomposite microspheres. Analysis of these microspheres showed complete incorporation of nanoparticles into the formulation. Optimal conditions for incorporation of nanoparticles into microspheres were using a composition of 75% nanoparticles and 25% lactose. The fine particle fraction of the microspheres was comparable to that of MF from the Asmanexregistered trademark TwisthalerTM. A biphasic release of MF was observed from the microspheres, with a significantly slower release compared to Asmanexregistered trademark. The spray drying process did not seem to alter the release properties of chitosan coated nanoparticles.
General Note: In the series University of Florida Digital Collections.
General Note: Includes vita.
Bibliography: Includes bibliographical references.
Source of Description: Description based on online resource; title from PDF title page.
Source of Description: This bibliographic record is available under the Creative Commons CC0 public domain dedication. The University of Florida Libraries, as creator of this bibliographic record, has waived all rights to it worldwide under copyright law, including all related and neighboring rights, to the extent allowed by law.
Statement of Responsibility: by Gina Patel.
Thesis: Thesis (Ph.D.)--University of Florida, 2010.
Local: Adviser: Hochhaus, Guenther.
Electronic Access: RESTRICTED TO UF STUDENTS, STAFF, FACULTY, AND ON-CAMPUS USE UNTIL 2011-08-31

Record Information

Source Institution: UFRGP
Rights Management: Applicable rights reserved.
Classification: lcc - LD1780 2010
System ID: UFE0041619:00001


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DEVELOPMENT OF NANOCOMPOSITE CORTICOSTEROID PARTICLES FOR USE
IN ASTHMA















By

GINA PATEL


A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL
OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT
OF THE REQUIREMENTS FOR THE DEGREE OF
DOCTOR OF PHILOSOPHY

UNIVERSITY OF FLORIDA

August 2010


































2010 Gina Patel



























This work is dedicated to my Mum, grandparents, husband and family, for their love and
support.









ACKNOWLEDGMENTS

I would like to express my sincere gratitude and appreciation to my mentor Dr.

Gunther Hochhaus, for accepting me into his group and his guidance and support

during the time I worked with him.

I would also like to thank my committee members, Dr. Hartmut Derendorf, Dr.

Anthony Palmieri III, Dr. Jeffrey Hughes and Dr. Christopher Batich for their guidance

on this project. I thank Dr. Arun Ranade for all the invaluable advice given to me in

guiding the development of the spray dryer device. During my studies I have been

fortunate enough to work as a teaching assistant for Dr. Cary Mobley, I would like to

thank him for making this time enjoyable and for always having the time to help and

advise on my research.

I would like to thank everyone in the Department of Pharmaceutics. I would like to

thank Yufei Tang for all the invaluable advice she gave me. I am grateful to all of the

past and present members of the Hochhaus group, Isabel Andueza, Vikram Arya, Intira

Coowantiwong, Navin Goyal, Manish Issar, Bhargava Kandala, Keerti Mudunuri, Elanor

Pinto, Srikumar Sahasranaman, Wan Sun, Nasha Wang, Benjamin Webber, Yanning

Wang and Kai Wu. For help with the experimental work over the years, I would like to

thank all of my former assistants, Christian Diestelhorst, Anica Liero, Gesa Nippel and

Pooja Patel. I must also thank Marc Rohrschneider for his advice and help on many

aspects of this project. I thank them all for their friendship, caring and support over the

years, for which I am truly grateful.

I would also like to show my gratitude and appreciation to Doug and Diane Ried

for welcoming me into their family while I participated in an internship in the College of

Pharmacy.









I would like to thank my Mum, grandparents, husband Jason Kwan and family for

all their love and support throughout my education. Without their encouragement, none

of this would have been possible.









TABLE OF CONTENTS



ACKNOWLEDGMENTS....................... ........... ..............................

L IS T O F T A B L E S .............................................................................. ....................... 8

L IS T O F F IG U R E S ....................................................................................................... 9

LIST O F A BBR EV IAT IO N S ... ... .................. .. ....................... ........................ 11

A B S T R A C T ....................................................................................................... ........... 14

CHAPTER

1 IN T R O D U C T IO N ........................................................................................................ 17

A s th m a ........................................................................................................ .......... 1 7
A sthm a T reatm ent ............................................................................................. 17
A diverse Effects of IC S ... .. .................... ................................................. 17
F a te o f IC S ............................................................................................................... .. 1 8
P ulm onary T targeting of IC S ............................................. ...... ............................. 19
Idea l C orticostero id ............................................................................................. 20
Inh a le r D ev ic e ................................................................................... .... ............ 2 1
Influence of Mass Median Aerodynamic Diameter.............................. ................. 23
C controlled Release Inhaled Form ulations............................................ ................ 23
Polym eric N anoparticles .......................................... ...................... .................... 25
Solid Lipid Nanoparticles (SLN) and Microparticles ..................... ................. 26
Low D ensity M icrospheres............................................................. ................ 26
P ulsed Laser D position (P LD ) ..................................................... ................ 27
O ligolactic A cid (O LA ) .......................................................................................28
P ro-drugs ........................................................................................................ 28
E ster Form ation .................................................................................................29
L ip o s o m e s ................... ... .............................................................................. 3 0
Enhancing M ucoadhesive Properties............................................................... 31
O b je c tiv e s ................................................................................................................... 3 1

2 DEVELOPMENT AND CHARACTERIZATION OF SLOW RELEASE
POLYMERIC CORTICOSTEROID NANOPARTICLES.........................................39

Introduction ............... ....................................................................... .... 39
H y p o th e s is .................................................................................................................. 4 2
M materials and M ethods ............................................................................................ 43
C h e m ic a ls .............................................................................................................4 3
Nanoparticle Preparation.......................................................................43
Drug loading and Encapsulation Efficiency......................................................45
Particle Size, Zeta Potential and Morphology ..................................................46


6









In vitro D rug R release S tudy ........................................................... ............... 46
R results a nd D discuss ion ............................................... .... ...... ...... ..... ........... ... 4 7
Influence of PLGA and PVA on Encapsulation Efficiency and Particle Size
of N a n o p a rtic les ..................................................................... .. ............... 4 7
Influence of PLGA and PVA on Morphology of Nanoparticles........................ 48
In Vitro Release Study of TA-PLGA Nanoparticles ....................................... 49
Chitosan Coated and Uncoated PLA MF Nanoparticles ................................ 50
C o n c lu s io n .............................................................................................................. 5 2

3 DEVELOPMENT OF NANOCOMPOSITE MICROSPHERES ............................... 71

In tro d u c tio n ............................................................................................................. 7 1
H y p o th e s is ............................................................................................................... 7 3
M materials and M ethods ........................................................................... 73
C h e m ic a ls ........................................... ........ ........................................................ 7 3
Development of Spray Dryer and Optimization of Operating Conditions ..........73
Spray Dried Chitosan Coated MF-PLA Nanoparticles..................................76
Nanoparticle Entrapment Efficiency and Loading into Nanocomposite
Microspheres .............................. ................. .............. 76
Morphology of Nanocomposite Microparticles............................................... 77
Determine MMAD of Nanocomposite Microparticles.................................... 77
In vitro drug R e lease S tudy ................................ ..................... ...................... 79
R es u lts a n d D isc uss ion ............... .................................. .. .... ................... .. ............ 7 9
Development of Spray Dryer and Optimization of Operating Conditions ..........79
Spray Dried Chitosan Coated MF-PLA Nanoparticles Influence of CH-
MF: lactose Ratio on Morphology of Spray Dried Microspheres ..................80
Spray Dried Chitosan Coated MF-PLA Nanoparticles Nanoparticle
In c o rp o ra tio n ....................... ............ .............. ................... .... ................ 8 1
Spray Dried Chitosan Coated MF-PLA Nanoparticles MMAD...................... 82
Spray Dried Chitosan Coated MF-PLA Nanoparticles In vitro release........... 83
C o n c lu s io n .............................................................................................................. 8 4

4 S U M M A R Y ............................................................................................................... 1 0 4

L IS T O F R E F E R E N C E S ........................................................................................... 106

B IO G RA PH ICA L S KETC H ......................................... ......................... ................ 117









LIST OF TABLES


Table page

1-1 PKPD properties of inhaled corticosteroids..................................................... 34

1-2 Average AUCs (n=3) in the lung, liver and brain and pulmonary targeting
(PT) in neonatal rats after intratracheal administration (50 pg/kg) of uncoated
budesonide and PLA coated budesonide........................................................ 36

1-3 Cumulative Receptor Occupancy (AUC), Pulmonary Targeting and Mean
Pulmonary Effect Times (MET) After Intratracheal Administration of
Escalating Doses of TAP in 800 nm Liposomes..............................................37

1-4 Influence of liposome composition on mucoadhesion and zeta potential ......... 38

2-1 Advantages and disadvantages of various methods for production of
n a n o p a rtic le s .......................................................................................................... 5 4

2-2 Particle size distribution for chitosan coated MF-PLA nanoparticles prepared
with 0.1% or 1% w/v chitosan by either incubation or in situ coating with
c h ito s a n ............................................................................................................... 6 6

2-3 Influence of chitosan coating to MF-PLA nanoparticles on encapsulation
efficiency, drug loading, particle size and zeta potential .................................67

3-1 MMAD cutoff for ACI analysis based on air flow rate (L/min)............................91









LIST OF FIGURES


Figure page

1-1 Fate of Inhaled C orticosteroids ........................................................ ................ 33

1-2 Pulsed laser deposition (PLD) Nanoclusters of polymer from the target are
deposited on larger micronized drug particles as continuous coatings that
sustain the release rate of drug in solution ..................................... ................ 35

2-1 Solvent evaporation technique to produce nanoparticles................................. 55

2-2 Schematic detailing methods to coat nanoparticles with chitosan .................... 56

2-3 Determine encapsulation efficiency and drug loading of nanoparticles ............ 57

2-4 Influence of PLGA and [PVA] on encapsulation efficiency of TA into
n a n o p a rtic le s .......................................................................................................... 5 8

2-5 Influence of PLGA and [PVA] on particle size of TA nanoparticles................... 59

2-6 SEM of TA-PLGA formulated with 200mg PLGA and 1% w/v PVA.................. 60

2-7 SEM of TA-PLGA formulated with 200mg PLGA and 2% w/v........................... 61

2-8 SEM of TA-PLGA formulated with 400mg PLGA and 1% w/v PVA.................. 62

2-9 In vitro release TA-PLGA nanoparticles formulated with 300mg PLGA using
1% o r 2 % w /v P V A ........................................................... ................ ............ 6 3

2-10 In vitro release of TA-PLGA nanoparticles formulated with 400mg PLGA and
1% o r 2 % w /v P V A ........................................................... ................ ............ 64

2-11 SEM of MF-PLA nanoparticles incubated with chitosan 1% w/v.......................65

2-12 SEM of uncoated MF-PLA nanoparticles formulated with 400mg PLA, 1% w/v
P V A ...................................................................................................... ....... .. 6 8

2-13 SEM of chitosan coated MF-PLA nanoparticles formulated with 400mg PLA,
1% w /v PVA and 0.1% w/v chitosan................................................ ................ 69

2-14 In vitro release of MF-PLA, chitosan coated MF-PLA (CH-MF) and
m ic ro n iz e d M F ........................................................................................................ 7 0

3-1 Spray dryer, designed to operate under lower temperatures than
com m ercially available instrum ents ................................................. ................ 86

3-3 Schematic to determine incorporation efficiency of nanoparticles into spray
d ried form u latio n ............................................................................................. 88









3 -4 H a n d iH a le r ........................................................................................................... 8 9

3-5 Cross section of Asmanex TwisthalerTM .......................................................... 90

3-6 MMAD of nanocomposite PLA-TA particles.....................................................92

3-7 Total PLA-TA nanoparticles for each MMAD range........................................ 93

3-8 Percentage of PLA-TA nanoparticles at each particle cutoff compared with
to ta l w e ig h t ........................................................................................................ 9 4

3-9 SEM of PLA-TA nanoparticles spray dried with lactose ................................. 95

3-10 Spray dried chitosan coated MF-PLA nanoparticles containing 10% of total
so lid feed as nanoparticles .............................................................. ................ 96

3-11 Spray dried chitosan coated MF-PLA nanoparticles containing 25% of total
so lid feed as nanoparticles .............................................................. ................ 97

3-12 Spray dried chitosan coated MF-PLA nanoparticles containing 50% of total
so lid feed as nanoparticles .............................................................. ................ 98

3-13 Spray dried chitosan coated MF-PLA nanoparticles containing 75% of total
so lid feed as nanoparticles .............................................................. ................ 99

3-14 Spray dried lactose, 1.25% w/v solid feed content ................. .................... 100

3-15 SEM of formulation contained in Asmanex TwisthalerTM .............................. 101

3-16 FPF of DD (%) of spray dried chitosan coated nanoparticles, to compare
batch to batch variability (n=3)...... ........ ..... ...................... 102

3-17 In vitro release of MF from spray dried CH-MF nanoparticles and Asmanex. 103









LIST OF ABBREVIATIONS

ACI Anderson cascade impactor

ACN Acetonitrile

AUC Area under the curve

AUCbrain Area under the receptor occupancy time profile of the brain

AUCliver under the receptor occupancy time profile of the liver

AUClung under the receptor occupancy time profile of the lung

BDP Beclomethasone dipropionate

BMP Beclomethasone monopropionate

BUD Budesonide

CFC Chlorofluorcarbon

CH-MF Chitosan coated mometasone furoate poly(lactic acid)
nanoparticles

CIC Ciclesonide

CIC-AM Ciclesonide active metabolite

CL Clearance

COPD Chronic obstructive pulmonary disease

DCM Dichloromethane

DDW Double distilled water

DD Delivered dose

Des-CIC Desisobutyryl ciclesonide

DPI Dry powder inhaler

DPPC 1,2-Dipalmitoyl-sn-Glycero-3-Phosphocholine

EE Encapsulation efficiency

F Bioavailability

FDA Federal drug administration









FLU Flutamide

Foral Oral bioavailability

FP Fluticasone propionate

FPF Fine particle fraction

fu Fraction of drug unbound

g Gram

GINA Global initiative for asthma

HFA Hydrofluroalkane

HPLC High performance liquid chromatography (Ultraviolet detection)

ICS Inhaled corticosteroid

kV Kilo volt

MAIC Major Analytical Instrument Center

MD Metered dose

MDI Metered dose inhaler

MeOH Methanol

MF Mometasone furoate

MF-PLA Mometasone furoate and poly(lactic acid) nanoparticles

mL Milliliter

ms Milisecond

MW Molecular weight

nm Nanometer

ODS Octadecyl silane

PBS Phosphate buffered saline

PLA Poly(lactic acid)

PLD Pulsed laser deposition









PLGA

pMDI

PT

PVA

RRA

RPM

SD

SEM

SMI

TA

TAP

TA-PLGA



TAP-lip

TAP-sol

Tg

Vdss

W

% v/v

% w/v

% w/w


Poly(lactic-co-glycolic acid)

Pressurized metered dose inhaler

Pulmonary targeting

Polyvinyl alcohol

Relative receptor affinity

Rotations per minute

Standard deviation

Scanning electron microscope

Soft Mist Inhaler

Triamcinolone acetonide

Triamcinolone acetonide phosphate

Triamcinolone acetonide and poly(lactic-co-glycolic acid)

nanoparticles

Triamcinolone acetonide phosphate liposomes

Triamcinolone acetonide phosphate solution

Glass transition temperature

Volume of distribution at steady state

Watt

% volume in volume

% weight in volume

% weight in weight









Abstract of Dissertation Presented to the Graduate School
of the University of Florida in Partial Fulfillment of the
Requirements for the Degree of Doctor of Philosophy

DEVELOPMENT OF NANOCOMPOSITE CORTICOSTEROID PARTICLES FOR USE
IN ASTHMA

By

Gina Patel

August 2010

Chair: GCnther Hochhaus
Major: Pharmaceutical Sciences

Asthma is a chronic inflammatory condition of the airways resulting in episodes

wheezing and breathlessness. Inhaled Corticosteroids (ICS) are used to prevent the

occurrence of asthma attacks. The clinical effect of ICSs depends on the time the drug

resides in the lung therefore increasing the drug residence time in the lung improves

asthma therapy. It has been proposed that nanoparticles could escape clearance

mechanisms in the lung and adhere strongly to the lung surface, leading to increased

residence time. There are two main barriers to this approach, firstly, nanoparticles

cannot deposit in the lung, and instead they are exhaled. Secondly, the particles must

be formulated to release drug slowly in order to take advantage of the increased

residence time.

In order to further improve lung targeting of corticosteroids, poly(lactic-co-glycolic

acid) PLGA and poly(lactic acid) PLA were used to produce polymeric nanoparticles of

Triamcinolone Acetonide (TA) and Mometasone Furoate (MF) using the solvent

evaporation technique. TA was used as a model corticosteroid to allow optimization of

various production parameters to produce PLGA nanoparticles. The solvent

evaporation technique was used to produce polymeric nanoparticles. A number of









process parameters may be varied to alter nanoparticle size and drug encapsulation

efficiency. Polyvinyl alcohol (PVA) surfactant concentration and PLGA content were

varied to determine their influence on particle size, drug encapsulation efficiency,

particle morphology and in vitro release characteristics of TA nanoparticles. Increasing

PLGA content resulted in a trend of increasing particle size and drug encapsulation. As

PVA concentration was increased particles tended to reduce in size and drug loading.

Nanoparticles produced ranged in particle size between 156-209nm. In addition when a

low concentration of 1 % or 2% w/v PVA was used to produce nanoparticles combined

with the use of only 200 mg PLGA, TA crystals were observed by scanning electron

microscopy. In vitro release studies revealed TA-PLGA nanoparticles released drug at

a similar rate to micronized TA, with 50% drug release being observed within 15

minutes. In order to produce nanoparticles which can deliver drug at a slower rate

compared to micron sized particles, a number of changes can be made to nanoparticle

production; a more lipophilic drug MF in combination with a more hydrophobic polymer,

PLA can be used to further slow down drug release. Subsequently MF nanoparticles

(MF-PLA) were using 10 mg MF, 400 mg PLA and 1% w/v PVA, these particles showed

slow release compared to MF contained in the Asmanex TwisthalerTM. To further

reduce MF release rate, nanoparticles were coated with chitosan. In vitro release

studies showed that chitosan coated MF-PLA nanoparticles (CH-MF) showed

significantly slower release compared to both uncoated nanoparticles and MF contained

within the Asmanex TwisthalerTM. In vitro release studies determined 100% MF

occurred after 1 hour for the Asmanex formulation, in comparison at this time only

50% and 24% MF was released from MF-PLA and CH-MF respectively.









A novel spray dryer was designed so that operating conditions of this device

allowed outlet temperatures to remain below the glass transition temperature of PLGA

and PLA, thus making this system suitable to spray dry polymeric nanoparticles.

Incorporation of these nanoparticles into lactose based microspheres by spray drying

resulted in spherical nanocomposite microspheres. Analysis of these microspheres

showed complete incorporation of nanoparticles into the formulation. Optimal

conditions for incorporation of nanoparticles into microspheres were using a

composition of 75% nanoparticles and 25% lactose. The fine particle fraction of the

microspheres was comparable to that of MF from the Asmanex TwisthalerTM. A

biphasic release of MF was observed from the microspheres, with a significantly slower

release compared to Asmanex. The spray drying process did not seem to alter the

release properties of chitosan coated nanoparticles.









CHAPTER 1
INTRODUCTION

Asthma

In the United States (US) asthma affects 16.4 million adults and 7 million children

and is increasing in prevalence [1-3]. Asthma can be defined as a complex

inflammatory disease of the airways involving many inflammatory mediators. Asthma is

a chronic reversible inflammatory disease of the airways that is characterized by

episodes of wheezing and breathlessness [4]. A component of asthma involves chronic

inflammation, if left untreated this can progress to airway remodeling, resulting in

irreversible airway narrowing [5].


Asthma Treatment

The Global Initiative for Asthma (GINA) was established in 1993 in order to focus

on reducing asthma prevalence, morbidity and mortality throughout the world.

Published guidelines for the treatment and prevention of asthma were produced by

GINA. Current therapy includes Inhaled Corticosteroids (ICS), long-acting P2 agonists,

leukotriene modifiers, theophylline and short acting P2 agonists for immediate relief [4].

ICS therapy targets the underlying inflammation present in asthma, to reduce long term

consequences such as progression of asthma to reversible obstruction in the airways.

ICS therapy is therefore the cornerstone for asthma treatment and is recommended for

all severity levels of persistent asthma by GINA [4, 6].


Adverse Effects of ICS

Pulmonary drug delivery has many attractive prospects for local and systemic

action. Selectively targeting drug delivery to the lung allows lower doses to be









administered providing a reduction in systemic side effects whilst maintaining clinical

efficacy. Although ICS are considered the "gold standard" in asthma management they

are responsible for many local and systemic adverse effects. Local side effects include

oropharyngeal candidiasis, dysphonia, perioral dermatitis (which may occur with the use

of spacer devices or nebulizers attached to face masks), cough, thirst and in rare cases

tongue hypertrophy. Additionally systemic side effects include adrenal insufficiency,

cataracts, glaucoma, growth retardation in children, osteoporosis, increased bone

fractures, skin thinning and skin bruising [7-11]. New developments in synthetic

corticosteroids have lead to improvements in pulmonary targeting and reductions in

adverse effects [12].


Fate of ICS

In order to produce further improvements in the area of ICS therapy, it is

necessary to be fully aware of the fate of corticosteroids following inhalation, as

described in Figure 1-1. Following inhalation a fraction of drug is deposited in the lungs,

a certain portion of the dose can be deposited in the oropharynx, and if this is not rinsed

may be swallowed. Subsequently the swallowed portion is available for oral absorption;

if the drug has significant oral bioavailability, it enters the systemic circulation to produce

unwanted adverse effects. Drug may enter the systemic circulation as a result of

absorption from the lung and gastrointestinal tract, both of which contribute towards

systemic availability of the ICS [12-14]. Newer ICS such as fluticasone propionate (FP)

and mometasone furoate (MF) have negligible oral bioavailability (Forai), therefore this

route does not significantly contribute to the systemically available dose [14]. A fraction

is deposited in the lungs where it must first undergo dissolution to exert its desired









pharmacologic effect; dissolved drug will be absorbed into the systemic circulation,

particulate matter deposited in the conducting airways however, will be subject to

clearance by the mucociliary escalator. In order to achieve high pulmonary targeting

drug dissolution should be slow, allowing pulmonary glucocorticoid receptors to be

occupied for a longer period of time. On the other hand if particles dissolve much

slower than the rate at which they are cleared from the lung, receptor occupancy over a

given time period will be reduced. In order to achieve pulmonary targeting,

corticosteroids should provide drug release at a rate comparable to their clearance from

the lung [15].


Pulmonary Targeting of ICS

Receptor occupancy-time profiles may serve as a surrogate marker for pulmonary

and systemic effects; cumulative receptor occupancy allows direct comparisons to be

made between local efficacy and systemic adverse effects [16]. For a given dose

increased cumulative receptor occupancy in the lungs compared to the systemic

circulation results in improved pulmonary targeting; providing better efficacy and

reduced systemic adverse effects [12, 14, 15]. Various factors are important for

pulmonary targeting of ICS, these are based on pulmonary or systemic factors.

Pulmonary pharmacokinetic factors include deposition efficiency and region of

deposition in the lung of the ICS, also pulmonary residence time and lung tissue binding

are important. Pulmonary pharmacodynamic factors include receptor binding affinity

and selectivity. Systemic factors include oral bioavailability, clearance, volume of

distribution, protein binding, tissue binding and affinity to transporters [15]. To









summarize, pulmonary targeting for inhaled drugs is influenced by both the inhaler

device and formulation and are discussed below [17].


Ideal Corticosteroid

An ideal ICS should have high pulmonary targeting, defined as the difference

between cumulative lung and liver receptor occupancies. High pulmonary versus

oropharnygeal deposition is desired, however as a significant fraction of inhaled drug is

deposited in the oropharynx from which it is subsequently swallowed, it follows that an

ideal ICS should posses negligible oral bioavailability. This is the case for newer

corticosteroids such as fluticasone propionate, mometasone furoate, ciclesonide and

des-ciclesonide, which have less than 1% oral bioavailability [18, 19]. Systemic

bioavailability is the sum of pulmonary and oral bioavailability, since oral bioavailability

is negligible for the newer corticosteroids, pulmonary absorption is responsible for the

significant portion of drug available to the systemic circulation. An ideal corticosteroid

should be removed rapidly from the systemic circulation to reduce systemic adverse

effects. The maximal clearance rate from the liver is approximately 90 L/min, which is

equal to the liver blood flow, most corticosteroids possess systemic clearance rates that

are similar to blood flow. ICS have high hepatic extraction, resulting in low systemic

exposure and enhanced pulmonary targeting, described in Table 1-1. Further

improvements in this area will be difficult as clearance of ICS are already close to that of

liver blood flow, improvements may be achieved through the possibility of increased

extra hepatic metabolism. Increased pulmonary residence prolongs drug action in the

lungs, thereby improving efficacy. Utilization of slow release inhaled formulations will

increase pulmonary residence time and targeting. In the case of slowly dissolving









drugs, the effect of mucociliary clearance predominates; as a result drug particles are

removed from the lungs before they are able to completely dissolve, thus reducing

pulmonary targeting [15]. Inhaled formulations capable of providing slow release

combined with reduced clearance from the lung are able to improve pulmonary

targeting; this approach increases efficacy allowing lower doses to be administered. As

a result fewer systemic side effects will be experienced.

Inhaler Device

Commonly used inhaler devices include Dry Powder Inhalers (DPIs) and

pressurized Metered Dose Inhalers (pMDIs) and to a lesser extent nebulizers, in

addition a new category of inhalers termed "soft" mist inhalers (SMI) has been

developed recently. pMDI devices utilize propellants to aerosolize a solution or

suspension, however they require the patient to co-ordinate both device actuation and

inspiration; resulting in poor inhaler technique, especially in young children and the

elderly. In addition to this a blast of high velocity cold propellant can impact the

oropharynx, in a small number of patients this will induce a gag reflex [20]. DPIs avoid

problems with inhaler technique as they are usually driven by the patient's own breath.

This itself however creates difficulty in many asthmatic and COPD patients who do not

have the necessary lung function [12].

The inhalation device plays a vital role in delivering drug to the lung. In the early

years of ICS therapy doses were delivered with low pulmonary deposition efficiencies of

approximately 10% or less [12]. In recent times however, new inhaler devices capable

of delivering >30% of the dose to the lung have been developed [13, 21, 22]. An

example of which comes from the re-development of the Qvar containing

beclomethasone dipropionate (BDP) from a CFC (CFC-BDP) to a HFA (HFA-BDP)









propellant, as a result pulmonary deposition increased from <10% to 50-60%. In

addition HFA-BDP deposited less drug on the oropharynx and resulted in a more even

distribution throughout the lung [21]. Modulite technology was developed to allow

HFA propellants to replace CFCs in pMDIs. The formoterol Modulite HFA inhaler

provides a respirable fraction of 35% and 32.5% respectively. Slower plume velocity

reduces impaction of the aerosol in the throat and therefore leads to greater lung

deposition. In addition slower plume velocities over a longer time period reduce

problems associated with poor hand-breath co-ordination; the inhaled dose is delivered

over a longer time period, therefore despite poor timing of the inhaler device with the

patient's inspiration, majority of the dose will still be inhaled. High speed photography of

the aerosol cloud produced following actuation of the Modulite device shows a greater

plume length, reduced velocity and extended spray duration of up to 220 ms with

formoterol Modulite compared to CFC propelled salmeterol. This approach leads to

reduced deposition in the oropharynx and subsequently improved pulmonary targeting

[23]. The Respimat belongs to a new class of inhaler devices termed "soft" mist

inhalers (SMIs). It does not contain any propellant, instead it utilizes the mechanical

energy from a spring to aerosolize droplets through a two channel nozzle, resulting in

the production of an aerosol as a result of impaction of two converging jets of liquid at a

carefully controlled angle. The uniblock is the key element of the Respimat,

constructed with a silicon wafer bonded to a small (2 mm x 2.5 mm) borosilicate glass

plate. The spray is generated over approximately 1.5 seconds; this results in a gentle

mist of respirable particles and allows more time for the patient to coordinate device

actuation with inspiration. As the Respimat produces a gentle mist which emerges at









a slower velocity compared to the aerosol cloud from pMDIS, less drug is deposited in

the oropharynx [22, 24, 25]. The Respimat delivers a greater fraction to the lung, it is

possible to reduce the Metered Dose (MD) and still maintain the same clinical efficacy.

The dose of ipratropium/fenoterol hydrobromide delivered using the Respimat may be

reduced by 50% of the dose administered using a pMDI with a spacer device, while

producing the same clinical effect [25]. Improvements in pulmonary deposition may

increase pulmonary selectivity, especially for corticosteroids with high oral

bioavailability. Improvements in lung deposition are not as significant for newer ICS

with low oral bioavailability [12].

Influence of Mass Median Aerodynamic Diameter

Deposition of particles in the lung is governed by its mass median aerodynamic

diameter (MMAD). Respirable particles are in the range 1-5pm and will deposit

effectively in the lung. Very small particles (<1 pm) will not be effectively deposited and

are exhaled. Large particles (>10pm) will be deposited in the tracheobronchial region

and then swallowed [26]. Depending on the oral bioavailability of the compound this

may contribute to therapeutic efficacy and adverse effects. For example many of the

more recently developed inhaled corticosteroids such as fluticasone propionate,

mometasone furoate, ciclesonide and des-ciclesonide have minimal oral bioavailability

thus this route does not contribute significantly towards systemic drug exposure [12].

Controlled Release Inhaled Formulations

A number of methods to prolong pulmonary residence time of inhaled formulations

have been employed. For example budesonide forms reversible fatty esters within

cells, this forms a depot of inactive drug until the ester is broken down to the active

form. This fatty acid esterification of budesonide, prolongs its action within the lungs









allowing once daily dosing and reduced systemic adverse effects [27-31]. In order for a

corticosteroid to undergo fatty acid esterification it must possess a steric-hindrance-free

hydroxyl group at the carbon 21 position [31]. Liposomal formulations have also been

investigated to provide slow release following inhalation, however, liposomes have

problems with stability due to leakage of drug during storage or jet-milling of lyophilized

formulations [16, 32]. Solid lipid nanoparticles and microparticles provide slow drug

release and increased stability in comparison to liposomes [33-35]. Biodegradable

polymers such as PLGA and PLA are used widely in the drug delivery due to their

biocompatible nature and approval for use as excipients [36]. PLGA microspheres have

been formulated to release dexamethasone continuously over one month [37]. Drug

particles may also be coated with nano thin layers of polymer to produce slow release

formulations, pulsed laser deposition of PLA onto glucocorticoids resulted in slower

release budesonide [38]. Large porous low density microparticles containing albuterol

demonstrated sustained bronchodilation over at least 16 hours compared to 5 hours

provided by non-porous particles of similar MMAD [39]. PLGA and PLA nanoparticles

have received a great deal of attention, a number of studies have described their use for

providing slow release. Usually a biphasic release profile was observed in most cases,

initially a burst release is observed, followed by a slow release profile [40-46]. Slow

release formulation of nanoparticles may be problematic, due to their large surface area

over which diffusion out of the polymer matrix occurs many studies observed very fast

drug release [47-49]. Mucoadhesive properties of chitosan allow nanoparticle retention

time to be further increased, a study by Yamamoto et al showed surface modified PLGA









nanoparticles, modified with chitosan had prolonged effects compared with unmodified

particles and also had reduced clearance[50].

Polymeric Nanoparticles

Polymer nanoparticles of PLGA or PLA are increasingly becoming the focus of

attention as they are biocompatible and biodegradable [36, 51]. Degradation products

are glycolic and lactic acid, natural bi-products of the Krebs cycle, readily eliminated

from the body by further breakdown to carbon dioxide and water [36, 51]. PLGA

degradation occurs by hydrolysis, and is dependent on molecular weight, conformation

and polymer composition [52, 53]. Rates of polymer degradation are fastest when the

composition is 50% lactic acid and 50% glycolic acid, thus we will use this for

nanoparticle production [53]. PLGA degradation however, may also be dependent on

the type of drug encapsulated [54].

Polymeric corticosteroid nanoparticles have been developed for a number of

medical applications, for example, cancer, arthritis, choroidal neovascularization and

neural drug delivery to name a few [40, 42, 43, 55-59]. Techniques for production of

polymeric nanoparticles include solvent evaporation, nanoprecipitation, supercritical

fluid precipitation, wet milling and high pressure homogenization [60-64].

The solvent evaporation technique is used commonly, it involves the production of

a microemulsion in which both drug and polymer are dissolved in the organic phase,

Figure 2-1 shows the scheme of production of nanoparticles using the solvent

evaporation technique. The organic phase is usually a volatile compound such as

dichloromethane or acetone, this will diffuse into the aqueous phase and evaporate,

leaving behind a nanoparticle suspension. Budhian et al explored the influence of

various production parameters on particle size and drug encapsulation efficiency [65].









Increased polymer lead to improved drug encapsulation, with a gradual increase in

particle size. PVA surfactant concentration less than 0.5% w/v resulted in a bimodal

particle size distribution due to insufficient stability of the microemulsion.

Solid Lipid Nanoparticles (SLN) and Microparticles

Solid lipid nanoparticles (SLN) have been studied for their potential in providing

sustained drug delivery [34]. Dexamethasone SLN particles have been developed

using soybean lecithin and glycerol tristearate. A biphasic in vitro release was observed

for these particles with an initial burst of approximately 70% followed by slower release

[35]. The lipid matrix of these SLN dispersion is more mobile in comparison to PLGA

nanoparticles, thus slow release drug formulations may be more challenging [34]. Slow

release salbutamol acetonide lipid microparticles have also been developed by Jaspart

et al. Encapsulation efficiency of the formulations developed was greater than 87%,

however high drug loading of around 25% w/w resulted in crystallization of drug

particles outside of the microparticles. In vitro release studies determined

microparticles formulated with lower drug loading produced a slower drug release,

however all solid lipid microparticles released at a slower rate compared to pure

salbutamol acetonide [33].

Low Density Microspheres

Large porous particles are more efficiently aerosolized as they produce fewer

aggregates and are easier to re-disperse within an air stream. It is also possible that

these large porous microparticles are too large to be engulfed by macrophages [66].

Large porous microparticles can be produced by spray drying, this results in particles

with a similar MMAD to smaller non-porous particles. Large porous estradiol particles

were aerosolized into the lungs of rats using an endotracheal tube. These particles









were shown to release over a longer time period of approximately 5 days compared to

only release over 1 day with non-porous particles of similar MMAD [67].

Pulsed Laser Deposition (PLD)

A nano-thin coating of a polymer such as PLA or PLGA may be applied with the

use of Pulsed Laser Deposition (PLD). High energy pulses of ultraviolet light are

directed onto a polymer disc to create a plume of nanoparticles which are subsequently

deposited onto a dry powder, thereby producing a nano-thin coating as shown in Figure

1-2 [68].

PLD can be used to produce sustained drug release, mean dissolution time for

budesonide particles coated was 4.7 + 0.1 hours compared to only 1.2 0.5 hours for

uncoated budesonide. Improved pulmonary targeting of PLA coated budesonide was

demonstrated in neonatal rats. Table 1-2 shows AUC of receptor occupancy-time

profiles of budesonide in neonatal rats in the lung and liver. It was seen that following

intratracheal administration of uncoated budesonide, AUC for the lung and liver were

indistinguishable from one another, whereas following administration of PLA coated

budesonide a higher AUClung was seen compared to AUCliver; indicating improvement in

pulmonary targeting with PLA coated budesonide [69].

Physicochemical differences between drugs also play an important role in

influencing improvements in pulmonary targeting following polymer coating.

Triamcinolone Acetonide (TA) and budesonide were coated with PLA, no difference

between dissolution profiles was observed between coated and uncoated TA, however

the more lipophilic corticosteroid, budesonide clearly showed slower drug release after

coating with PLA [70].









Oligolactic Acid (OLA)

Oligolactic acids (OLAs) are short chain versions of polylactides, which are

biocompatible and approved by the FDA for use with implantable devices. OLAs are

formulated as excipients in MDI inhalers by 3M, they have been used as solubilizers,

suspending agents and produce sustained release within the lung. Sustained release is

achieved by the formation of OLA-drug matrix (solution formulations) or OLA-coated

drug particle (suspension formulations) [71, 72].

Pro-drugs

Many ICS, such as BUD or FP are administered in their pharmacologically active

form; others may be inhaled as pro-drugs that must first undergo conversion to their

active form. This approach can be utilized to improve pulmonary by achieving

therapeutic drug concentrations at the target site whilst minimizing unwanted side

effects at other sites. Deposition of ICS in the mouth and oropharynx can give rise to

adverse effects such as oral candidiasis and dysphonia. Two currently available ICS

include beclomethasone dipropionate and ciclesonide are converted to their active

metabolites by esterase enzymes within the pulmonary epithelium [73-75]. It has been

shown that bioactivation of ciclesonide is very low in the oropharynx, therefore less

active drug is present in comparison administration of budesonide or FP [14].

Ciclesonide, a newer inhaled corticosteroid is converted to its active metabolite by ester

cleavage at the C21 position, resulting in the formation of desisobutyryl-ciclesonide with

100-fold higher potency. Combined with high systemic clearance of both the parent

compound and the active metabolite, local and systemic adverse effects are reduced.

In addition the active metabolite of ciclesonide undergoes fatty acid esterification within

the lung, to prolong retention within the lung, described in further detail later [74-77].









Ester Formation

Pulmonary retention of inhaled corticosteroids may be enhanced due to reversible

fatty acid esterification. Conjugation has only been reported for budesonide,

triamcinolone acetonide and des-ciclesonide. This process occurs within cells in the

lung and forms a reservoir of lipid conjugated drug. Conjugated corticosteroid is slowly

hydrolyzed by enzymes to release free drug. In order for this fatty acid conjugation to

take place the ICS must possess a steric-hindrance-free hydroxyl group at the carbon

21 position. Reversible fatty acid conjugation provides increased anti-inflammatory

action at the target site resulting in improved pulmonary selectivity [27-31, 74, 75, 78-

80]. In general, rapid ester formation and slow ester hydrolysis leads to improved

pulmonary targeting; this is further improved in combination with a high systemic

clearance.

The corticosteroid budesonide is moderately lipophilic with a relatively fast

dissolution, followed by a rapid absorption from the lung. Budesonide would not be

expected to have a prolonged duration of action in the lung when compared to

fluticasone propionate, a corticosteroid with both higher lipophilicity and relative receptor

affinity [80]. Formation of the fatty acid ester, budesonide oleate however, prolongs

retention in the airways and allows for once daily dosing [29].

Ciclesonide is a corticosteroid which first undergoes metabolism to the active form,

des-ciclesonide followed by reversible fatty acid ester formation. This ester formation

was confirmed to occur in the human lung in a single dose, open-label, non-randomized

study in 20 patients. The metabolites des-ciclesonide, des-ciclesonide-oleate and des-

ciclesonide-palmitate were detected in the central and peripheral lung tissue [73].









Budesonide, however was shown to be esterified more rapidly and to a greater extent in

comparison to ciclesonide in rat tracheal tissue [79].

Future developments in corticosteroids may involve selection of compounds which

undergo fatty acid esterification [81].

Liposomes

Liposomes have been extensively studied due to their ability to incorporate both

hydrophilic and hydrophobic drugs, as well as ability to produce a variety of particle

sizes and have been used as drug carriers since the late 1960s [82]. Liposomes are

composed of phospholipids which are endogenous in the lung, improving compatibility.

Nebulized BDP-DLPC liposomes administered to healthy volunteers are well tolerated

in doses equivalent to those currently used for treatment in asthma [83]. Modified

liposomes may be used to target delivery to different cells, liposomes prepared with

mannosylated cholesterol derivatives can enhance uptake into macrophage cells [84].

Liposomal formulations have also been investigated to provide slow release following

inhalation, however, liposomes have problems with stability due to leakage of drug

during storage or jet-milling of lyophilized formulations [16, 32]. Solid lipid nanoparticles

and microparticles provide slow drug release and increased stability in comparison to

liposomes [33-35]. It has been demonstrated that slow drug release from liposomes is

able to improve pulmonary selectivity in the rat model. Triamcinolone Acetonide

Phosphate (TAP) liposomes sized 200nm and 800nm were shown to produce biphasic

drug release in vitro, with the later resulting in slower release). In vivo studies in male

F-344 rats were performed in order to compare pulmonary selectivity of these liposomes

with TA solution and liposomes. TA liposomes release drug rapidly under sink

conditions, and thus both the TA solution and liposome formulation would not be









expected to produce pulmonary targeting, Table 1-3 show 800nm TAP liposomes

produced the greatest pulmonary targeting; pulmonary targeting is defined as the area

under the curve of the receptor occupancy time profile of the lung compared to the liver

(Pulmonary targeting (%*hr) = AUClung AUCliver). Slow drug release from liposomes

allows reduction in dose and dosing frequency, thereby reducing systemic adverse

effects observed whilst maintaining clinical effects in the lung [16]. It has been shown

that budesonide encapsulated into stealth liposomes, delivered once a week was able

to provide an equivalent anti-inflammatory effect as once daily administration of

budesonide [85].

Enhancing Mucoadhesive Properties

Chitosan is a polysaccharide containing an amino group which may be positively

charged, particles composed or coated with chitosan increase interaction with

negatively charged lung epithelial cells. In addition chitosan may also influence release

characteristics of drug particles. Gelatin has also been utilized to produce particles.

Rifampicin liposomes coated with chitosan had greater mucoadhesive properties and

lower toxicity towards A549 epithelial cells compared to uncoated liposomes. Table 1-4

demonstrates the relationship between zeta potential and mucoadhesive properties of

rifampicin liposomes. Uncoated negatively charged liposomes showed the lowest

amount of adhesion, followed by uncharged liposomes, with positively charged chitosan

coated liposomes showing the most mucoadhesion [86].

Objectives

Increased pulmonary retention has been observed following inhalation of

nanometer sized particles in comparison to micronized particles [87, 88]. In addition

polymers such as PLGA and PLA are able to reduce drug release from nanoparticles.









We hypothesize that nanoparticles may be spray dried with lactose to produce

nanocomposite microspheres; these microspheres will be capable of delivery to the lung

following inhalation. The above hypotheses will be tested by the following specific aims;



Preparation and characterization of slow release polymeric corticosteroid
nanoparticles.

In vitro drug release testing to determine slow release characteristics of
nanoparticles.

Design of a spray dryer able to enable nanoparticles to be spray dried at
temperatures lower than operating temperatures of commercially available spray
dryers.

Preparation of spray dried nanoparticles to form nanocomposite microspheres
with MMAD in the respirable range.

Determine in vitro release characteristics of spray dried formulations to
investigate the effect of spray drying on release rate.











Lung f
d sition ATM from luI

Mucociliary
clearance "-
\ Orally
Swallowed absorbed
kfrUtion fraction

|I fl-t LUV

Absorption
from gut


First-pass
inactivation

Figure 1-1. Fate of Inhaled Corticosteroids [12]


SYSTEMIC
SIDE
EFFECTS









Table 1-1. PKPD properties of inhaled corticosteroids [15]
ICS RRA CL (L/hr) Vdss (L) Forai (%) fu (%)

BDP 53 150 20 15-20 13

BMP 1022 120* 424 26 -

FLU 190 57 96 20 20

TA 233 37 103 23 29

BUD 935 84 18,311 12

FP 1800 69 318 <1 10

MF 2900 54 <1 1-2

CIC 12 140 207 <1 1

Des-CIC 1200 228* 897 <1 <1


























Figure 1-2. Pulsed laser deposition (PLD) Nanoclusters of polymer from the target are
deposited on larger micronized drug particles as continuous coatings that
sustain the release rate of drug in solution [38]









Table 1-2. Average AUCs (n=3) in the lung, liver and brain and pulmonary targeting
(PT) in neonatal rats after intratracheal administration (50 pg/kg) of uncoated
budesonide and PLA coated budesonide [69]
Formulation Dose AUClung AUCliver AUCbrain PT


(pg/kg)


(AUClung/AUCliver)


Uncoated 50 58.4+ 12.9 56.4 +6.8 38.3+6.7 1.03+0.13

budesonide

PLA coated 50 75.8 +3.7 46.6 + 14.5 29 + 7 1.72 +0.46

budesonide











Table 1-3. Cumulative Receptor Occupancy (AUC), Pulmonary Targeting and Mean
Pulmonary Effect Times (MET) After Intratracheal Administration of
Escalating Doses of TAP in 800 nm Liposomes [16]


TAP-sol


TA-lip 200


TAP-lip 200


TAP-lip 800


Lung

Liver


Pulmonary Targeting

(%*hr)(AUClung-

AUCliver)

Pulmonary Targeting

(%*hr)(AUClung/AUC iver)

Mean Pulmonary Effect

Time (hr)


0.85


>6.2


370 + 50

340 + 40

30 + 10


320 + 85

380 + 10

-60 + 80


770 + 120

620 + 150

150 + 60


1070 + 70

700 + 140

370 + 70









Table 1-4. Influence of liposome composition on mucoadhesion and zeta potential [86]
Liposome composition Mucin adsorbed on Zeta-Potential (mV)
liposomes (%) Mean
value (SD)
PC/Chol 17.0+8.3 +0.09+0.54


PC/PG/Chol
[PC/Chol]cHT
[PC/PG/Chol]cHT
DSPC/Chol
DSPC/PG/Chol (9:1:5)
[DSPC/Chol]cHT
[DSCP/PG/Chol]cHT


7.4 + 4.4
47.1 + 1.2
90.9 + 7.6
46.2 +4.1
25.1 +8.1
66.6 + 2.2
93.1 +4.1


-22.9 +2.1
+4.4+1.9
+24.98 + 0.91
+0.93 + 0.77
-19.9 +2.3
+5.4 + 2.7
24.43 + 0.62









CHAPTER 2
DEVELOPMENT AND CHARACTERIZATION OF SLOW RELEASE POLYMERIC
CORTICOSTEROID NANOPARTICLES

Introduction

Nanoparticles may have the possibility of providing increased retention in the

airways compared to micron sized particles [87]. Nanoparticles have been extensively

studied in the field of toxicology. A number of studies have been conducted in order to

determine deposition, retention and translocation of ultrafine particles in the lung.

Technetium Tc 99m (99mTc)-radiolabeled 100nm carbon particles were administered to

healthy subjects and COPD patients by nebulizer. The central/peripheral (C/P)

distribution was controlled by administering either a shallow or deep aerosol bolus. A

shallow aerosol bolus is used to deliver either more centrally compared to a deep

aerosol bolus inhalation. 48 hours following a shallow aerosol bolus, 70% and 82% of

particles were retained within the airways of healthy non-smokers and COPD patients

respectively. It has been shown that nanoparticles are retained in the lung for a longer

time period compared with micron sized particles; 24 hours following inhalation >70% of

nanoparticles are retained in the airways, in comparison only 10% of particles greater

than 6tm are retained [87, 88]. It has been proposed that increased retention on

nanoparticles in the lung is as a result of greater displacement of nanoparticles into the

aqueous surfactant film compared to micron sized particles [87, 89, 90].

In recent times nanotechnology has received much attention, potential applications

include imaging and diagnosis, targeted drug delivery and controlled drug release [91,

92]. Nanotechnology itself is not a new concept, liposomes were first used as drug

carriers in the late 1960s [82]. Polymer nanoparticles of PLGA or PLA are increasingly

becoming the focus of attention as they are biocompatible, biodegradable and have









been approved by the FDA for use in implantable devices [36, 51]. Degradation

products are glycolic and lactic acid, natural bi-products of the Krebs cycle, readily

eliminated from the body by further breakdown to carbon dioxide and water [36, 51].

PLGA degradation occurs by hydrolysis, and is dependent on molecular weight,

conformation and polymer composition [52, 53]. Rates of polymer degradation are

fastest when the composition is 50% lactic acid and 50% glycolic acid, thus this was

used for production of nanoparticles [53]. PLGA degradation however, may also be

dependent on the type of drug encapsulated [54]. Drug release from particles is also

influenced by polymer composition. Increased lactide:glycolide results in slower

release, possibly due to increased hydrophobicity of the polymer as well as increased

solid state solubility of hydrophobic drugs in the polymer matrix. It has also been

determined that PLGA with an ester terminated end group released hydrophobic drugs

at a slower release rate compared with a carboxylic acid end group [93].

Polymeric corticosteroid nanoparticles have been developed for a number of

medical applications, for example, cancer, arthritis, choroidal neovascularization and

neural drug delivery to name a few [40, 42, 43, 55-59]. Techniques for production of

polymeric nanoparticles include solvent evaporation, nanoprecipitation, supercritical

fluid precipitation, wet milling and high pressure homogenization [60-64]. Advantages

and disadvantages of the various methods for nanoparticle production are discussed in

Table 2-1.

The solvent evaporation technique is used in the production of monodispersed

spherical polymeric nanoparticles. This technique is used commonly, it involves the

production of a microemulsion in which both drug and polymer are dissolved in the









organic phase, Figure 2-1 shows the scheme of production of nanoparticles using the

solvent evaporation technique. The organic phase is usually a volatile compound such

as dichloromethane or acetone, this will diffuse into the aqueous phase and evaporate,

leaving behind a nanoparticle suspension. Nanoparticles may be collected and washed

by centrifugation, followed by lyophilization to allow storage as a dry powder. A number

of process parameters may be manipulated to influence particle size and drug loading.

Budhian et al explored the influence of various production parameters on particle size

and drug encapsulation efficiency [65]. Increased polymer content lead to improved

drug encapsulation, with a gradual increase in particle size. Increased polymer content

allows more drug to be dispersed within the polymer matrix, thus resulting in higher

encapsulation efficiency. As the polymer contained in the organic phase increases

however, viscosity of this phase is also increased, thus resulting in formation of larger

o/w microemulsion droplets and larger nanoparticle size. PVA surfactant concentration

less than 0.5% w/v resulted in a bimodal particle size distribution due to insufficient

surfactant concentration for production of a stable microemulsion.

The solvent evaporation method has been used to produce slow release

nanoparticles of a number of different compounds including paclitaxel, etanidazole,

bezopsoralen, flurbiprofen and dexamethasone; a biphasic in vitro release was

observed in most cases [40-46]. It is difficult to compare results from different studies

however, as drug release rates depend on release conditions such as presence of sink

conditions, release media and sample separation methods (filtration, centrifugation or

dialysis). Though each method has its advantages it is possible that certain artifacts

may result due to the separation method used. For example retardation of drug release









from dialysis bags may cause release profiles to reflect diffusion through the dialysis

membrane rather than from the formulation [34, 94]. In these situations, it is possible

for the drug to interact with the dialysis membrane, slow release observed is due to an

interaction with the membrane rather than the actual release rate from the nanoparticle

formulation; therefore it is necessary to ensure that the chosen dialysis membrane does

not interact with the drug of interest, as well as inclusion of adequate controls in the

experiment. In addition, release rate can further be reduced by coating with another

polymer such as chitosan. Chitosan is polysaccharide derived from the deacetylation of

chitin, commonly obtained from the shells of crustaceans. It is biocompatible and can

be degraded by lysozymes present in all mammalian cells [95]. Cationic chitosan is

able to coat nanoparticles due to an electrostatic interaction with the negatively charged

carboxylic acid end group present on the polymer PLGA or PLA [96]. Chitosan has

been used to coat PLGA nanoparticles as well as liposomes [86, 96-99]. PLGA

nanoparticles surface modified with chitosan also had increased retention in the lung

compared to PLGA nanoparticles due to improved mucoadhesive properties of chitosan

[50].

Hypothesis

We hypothesize that polymeric nanoparticles with a monodispersed particle size

distribution can be formed using the solvent evaporation method. Various parameters

during this process will be manipulated to influence particle size and drug

encapsulation. We hypothesize that increasing polymer content will increase particle

size and drug encapsulation efficiency. Also increasing PVA concentration will allow

smaller oil-in-water microemulsion droplets to be formed, resulting in smaller particles

with lower drug encapsulation. Larger particles may be formed with reducing PVA









concentration, however a minimum content must be achieved in order to sufficiently

stabilize the microemulsion and result in a monodispersed particle size distribution.

It is hypothesized that chitosan coated nanoparticles will be able to release at a

slower rate in comparison with uncoated particles. Electrostatic interaction between the

positively charged chitosan and negatively charged polymer allows chitosan to coat

polymeric nanoparticles. It is important that the polymer chain is terminated with a

carboxylic acid end group and is not ester terminated, to ensure interaction with the

chitosan. Methods to produce chitosan coated particles will be investigated and

optimized.

In vitro release testing will be performed by the batch/filter method [100]. This

method reduces artifacts which may be observed by using the dialysis method for

separation of particles from free drug. It is hypothesized that polymeric nanoparticles

will release drug at a slower rate compared to the control micronized drug.


Materials and Methods

Chemicals

Micronized TA was purchased from PCCA Inc. (Houston, TX, USA). MF was

donated by Ipca laboratories Itd (Mumbai, India). PLGA and PLA were purchased from

Lactel Absorbable Polymers (Pelham, AL, USA). DCM and ACN were purchased from

Fisher Scientific (Pittsburgh, PA, USA). PVA and medium molecular weight chitosan

were obtained from Sigma Chemical Co. (St. Louis, MO).

Nanoparticle Preparation

Polymeric nanoparticles of either TA or MF were prepared using the solvent

evaporation technique [47]. TA will be used as a model drug to allow investigation of









process parameters on particle size, drug encapsulation efficiency and in vitro release

profiles. The polymer used was either DL-PLGA (50:50 inherent (0.55-0.75 dL/g

inherent viscosity) or DL-PLA (0.54 dL/g inherent viscosity). In brief, 10 mg of drug and

200-400 mg polymer were dissolved in 5 mL DCM, this was pre-emulsified with 5 mL of

the aqueous PVA phase by vortexing for 30 seconds in a 20 mL glass scintillation vial

using the Fisher vortex genie 2 at speed 9. The pre-emulsion was added to the

remaining 45 mL PVA solution and sonicated on ice (Sonics Vibra Cell Ultrasonicator

Newtown, CT) at 60 W for 5 minutes to form an oil-in-water microemulsion. PVA

concentration was varied between 1 and 3% w/v for each PLGA level, (n=3). DCM was

allowed to evaporate under gentle stirring on the magnetic stirrer for 4 hours to produce

a nanoparticle suspension which was collected by centrifugation at 20,000 rpm for 40

minutes and washed Beckman J2-21 (Beckman Coulter, Inc.Fullerton, CA) using the

JA-20 rotor. Nanoparticles were lypophilized using a Labconco freeze dryer (Labconco

Corporation, Kansas City, MO). Dried nanoparticles were stored in amber glass vials at

40C in a dessicator.

Two different methods to prepare chitosan coated PLA nanoparticles were

studied. Figure 2-2 shows details of these two methods, the incubation method involves

dispersion of PLA nanoparticles in either 0.1% or 1% w/v chitosan solution. In situ

coating allows the chitosan coating to be applied in the same step as formation of the

nanoparticles. Nanoparticles are prepared using the solvent evaporation technique,

however chitosan is already dissolved in the aqueous PVA phase of the microemulsion.

During formation of the nanoparticle suspension, the positively charged chitosan is

coated onto the negative PLA particles by an electrostatic interaction. The optimal









method was chosen based on the ability to form unimodal nanoparticles. The aqueous

phase of the oil-in-water microemulsion was produced using PVA and chitosan in 1%

v/v acetic acid in DDW. 400 mg PLA and 10 mg MF were dissolved into 5 mL DCM,

this was pre-emulsified with 5 mL of the aqueous phase by vortexing for 30 seconds.

This was combined with the remaining 45 mL of the aqueous phase and sonicated on

ice for 5 minutes at 60 W to produce an oil-in-water microemulsion. The emulsion was

gently stirred for a further 4 hours to allow the DCM to evaporate and leave behind a

chitosan coated nanoparticle suspension. The resultant suspension was centrifuged

and washed using a 1 % v/v acetic acid solution. Chitosan coated nanoparticles were

dispersed in a small volume of double distilled water followed by lyophilization and

stored in amber glass vials in a desiccator at 40C.

Drug loading and Encapsulation Efficiency

In order to determine the drug content of the lyophilized nanoparticles

approximately 5 mg of the dry powder was weighed. The particles were then dissolved

using 2 mL DCM and placed on an orbital shaker (Bellco biotechnology, Vineland, NJ)

overnight to ensure disintegration of the nanoparticles. The DCM was evaporated off

using a Jouan RC10.10 vacuum centrifuge (Thermal Fisher, Asheville, NC), the dried

residue was dissolved in 1 mL mobile phase (ACN:DDW, 60:40) and analyzed by

HPLC-UV (Hewlett Packard Series 1050) using a Phenomenex Ultracarb 30 4.6x150

mm ODS column using mobile phase at a flow rate of 1.2 mL/min sample peaks were

quantified at 254 nm wavelength. Concentrations for the caliberation curve were 10, 20,

40, 60, 80 and 100 [tg/mL with an R squared value of at least 0.997. Figure 2-3

describes the method by which EE and drug loading of nanoparticles was determined.









Drug loading and encapsulation efficiency were determined by the equations below.

Theoretical drug loading (% w/w) was calculated by calculating how much drug would

be theoretically present in 100 g of the lyophilized powder provided that no drug or

polymer is lost during the nanoparticle production process, the equation for which is

described below.

Amount of drug contained in NP
Actual Drug loading (% w/w) = x 100
Total weight of NPs weighed

Drug weighed
Theoretical drug loading (% w/w) = x 100
Drug + Polymer weighed

Actual drug loading (% w/w)
Encapsulation Efficiency (% w/w) = x 100
Theoretical drug loading (% w/w)

Particle Size, Zeta Potential and Morphology

Particle size and zeta potential were determined using the Nanotrac (Microtrac)

and Brookhaven ZetaPlus. SEM was performed to observe particle morphology using

scanning electron microscope (SEM) JEOL JSM-6335F instrument (Major Analytical

Instrument Center (MAIC), UF, Gainesville, FL). Briefly, formulations were placed on

carbon stubs which were coated with carbon using a vacuum evaporator. SEM was

conducted using 2 kV.

In vitro Drug Release Study

In vitro release was performed in 100 mL 1% v/v tween 80 in PBS, shaken at 30

rpm in a hot shaker (Bellco biotechnology, Vineland, NJ) at 370C over 24 hours under

constant sink conditions. Tween 80 was used to increase the saturation concentration of

TA and MF in the dissolution media to 83 pg/ml and 25 pg/ml with addition of 1% v/v

tween 80 at 370C. In order to maintain sink conditions maximal concentrations obtained

did not exceed 10% of the saturation concentration. At specific time points (0, 15, 30,









90, 120, 180, 240, 360, 600 and 1440 minutes), 1 mL was removed using a pipette and

subsequently filtered using a 0.02 |tm filter (Whatman Anotop plus filter), 1 mL fresh

buffer was replaced into the dissolution media. The maximum concentration obtained

was determined based on analysis of remaining undissolved drug at the end of the

dissolution testing. After the last time point, the dissolution media was centrifuged at

20,000 rpm for 40 mins using the Beckman J2-21 (Beckman Coulter, Inc.Fullerton, CA)

using the JA-20 rotor. The pellet was dissolved in 2 mL DCM and placed on an orbital

shaker overnight (Bellco biotechnology, Vineland, NJ) to ensure disintegration of the

particles. HPLC was used to quantify released drug concentrations using a caliberation

curve ranging from 0.5-10 ug/mL. Controls used were micronized TA and MF contained

in the Asmanex formulation from the reservoir based inhaler device, TwisthalerTM,

Schering-Plough.

Results and Discussion

Influence of PLGA and PVA on Encapsulation Efficiency and Particle Size of
Nanoparticles

A factorial design was implemented to determine influence of PLGA and PVA

concentration on particle size and encapsulation of TA within nanoparticles.

Encapsulation efficiency of TA nanoparticles decreased with increasing PVA

concentration, shown in Figure 2-4 as expected. Increased PVA surfactant

concentration in the microemulsion aqueous phase results in an increased solubility of

TA in the aqueous phase. A consequence of this is increased loss of drug during

nanoparticle production is reduced encapsulation efficiency and drug loading.

Increasing PLGA content did not have any significant effect on encapsulation efficiency.

Increasing PVA concentrations resulted in a trend of decreased particle size, as shown









in Figure 2-5. Increased surfactant concentration in the oil-in-water microemulsion

leads to the formation of smaller oil phase droplets; as a result particle size of

nanoparticles decreases with increased PVA concentration. Increasing PLGA content

results in the formation of larger nanoparticles, this is as a result of increased viscosity

of the oil phase of the microemulsion; larger organic phase droplets are formed leading

to an increase in particle size. This is in agreement with studies performed by Budhian

et al; increasing PVA concentration results in reduced drug loading initially, between 1-

2% w/v PVA then plateau. Particle size also decreased with increasing PVA

concentration in their study, however with the use of >5% w/v PVA particle size

increased. This is due to competing effects of increased stabilization of the

microemulsion with increasing surfactant, which reduces particle size. High PVA

concentration increases viscosity of the aqueous phase, leading to reduced net shear

stress for droplet breakdown during formation of the microemulsion, therefore resulting

in larger particles [65]. The PVA concentration range we investigated was 1-3% w/v,

thus it was determined increasing PVA concentration reduced encapsulation efficiency

and a trend of reduced particle size was observed. Budhian et al also investigated the

influence of PLGA content on particle size and drug loading. Increasing PLGA content

lead to a gradual increase in particle size and drug loading, as was observed in our

experiments producing TA-PLGA nanoparticles.


Influence of PLGA and PVA on Morphology of Nanoparticles

TA crystals were observed for two formulations of TA-PLGA nanoparticles, these

were both formulated using 200mg PLGA using 1% and 2% w/v PVA, SEM of these

formulations are shown in Figure 2.6 and Figure 2.7. These TA crystals were only









observed when particles were formulated with both low polymer contents and surfactant

concentrations. There must be sufficient polymer for the drug to disperse within in the

solid state; drug not present in the polymer matrix must then be washed off [93]. When

only 200 mg PLGA is used, there is insufficient polymer for the TA to disperse within,

coupled with low PVA surfactant, unencapsulated drug is not washed away leaving

behind free TA crystals. Although it appears that encapsulation efficiency for these two

samples is greater than 50%, this is not due to TA trapped in PLGA nanoparticle matrix.

Figure 2.8 shows spherical TA-PLGA nanoparticles produced with 400 mg PLGA and

1% w/v PVA, no TA crystals were observed in this formulation. Formulations produced

using 300-400 mg PLGA did not contain any unencapsulated drug crystals, as

determined by SEM.

In Vitro Release Study of TA-PLGA Nanoparticles

Various methods to determine in vitro drug release of nanoparticles have been

reported in the literature. Regenerated cellulose ester membranes and dialysis bags

have been used to physically separate nanoparticles from the dissolution media.

Preliminary studies performed however, indicated interaction of corticosteroids with the

membrane and delayed release of TA from within dialysis bags. If this method is used,

it is essential to find dialysis membranes which do not interact with the drug being

studied, as well as use of adequate controls [94]. Separation by filtration was used to

determine in vitro release to avoid problems which occurred with the use of regenerated

cellulose dialysis membranes.

In vitro release studies were performed on TA-PLGA nanoparticles formulated

with either 300 mg or 400 mg PLGA using both 1% and 2% w/v PVA. These

formulations had good drug encapsulation efficiency as well as a unimodal particle size.









In comparison to the control of micronized TA, nanoparticle formulations made with only

300mg PLGA released at a slightly faster rate, as shown in Figure 2-9. TA-PLGA

nanoparticles formulated with 400mg PLGA were shown to release at the same rate as

micronized TA, refer to Figure 2-10. All formulations tested release drug at a very

similar rate to the control micronized TA. The surface area to mass ratio of

nanoparticles is very high compared with micron sized particles, thus if no modification

is made to the formulation, nanoparticles would release drug at a much faster rate

compared to their micron sized counterpart. TA-PLGA nanoparticles posses some slow

release characteristics as they are able to release drug at a similar rate to micronized

TA.

Chitosan Coated and Uncoated PLA MF Nanoparticles

Two methods for production of chitosan coated polymeric nanoparticles were

investigated. The chitosan coating procedure was optimized to produce particles

coated particles with a unimodal particle size distribution. Lyophilized MF-PLA

nanoparticles that were incubated with either 0.1 % or 1 % w/v chitosan solutions

resulted in nanoparticles that had multimodal particle size distributions. A possible

explanation for this is due to incomplete dispersion of nanoparticles in the viscous

chitosan solution, leading to coating of agglomerated nanoparticles. SEM of MF-PLA

nanoparticles incubated with a chitosan solution showed a bimodal particle size

distribution (Figure 2-11). In situ chitosan coating involves coating during the formation

of nanoparticles. Chitosan is present in the aqueous phase of the microemulsion,

positively charged chitosan adheres onto the negatively charged PLA nanoparticles.

Addition of 1 % w/v chitosan to the aqueous phase results in bimodal particle size

distribution. High viscosity of 1% w/v chitosan solution resists droplet breakdown during









sonication, producing some larger particles. In addition, particles produced from 1% w/v

chitosan solutions are more difficult to collect by centrifugation, due to the high viscosity

of the solution. The use of 0.1% w/v chitosan within the situ coating procedure resulted

in a unimodal particle distribution and can be seen by SEM (Figure 2-13) and can be

compared to uncoated MF-PLA nanoparticles (Figure 2-12). Particle size of chitosan

coated particles was 226 nm compared to 212 nm for uncoated particles. Although

particle size does not significantly increase following coating with chitosan, the presence

of chitosan is confirmed by the reversal of zeta potential from negative charge on

uncoated nanoparticles to positive charge on coated particles, as shown in Table 2-3.

In order for chitosan to coat particles, the PLA polymer must terminate with a carboxylic

acid group and the amine group on chitosan must be protonated, with 1 % v/v acetic

acid.

In vitro release of chitosan coated particles was compared to uncoated MF-PLA

nanoparticles and Asmanex as a control are shown in Figure 2-14. A biphasic release

was observed for both the chitosan coated and uncoated nanoparticles. Initially the

burst release of MF-PLA nanoparticles is similar to the Asmanex formulation, followed

by slow release. Chitosan coated MF-PLA nanoparticles (CH-MF) release

approximately 20% by 30 minutes, followed by a much slower release rate compared to

MF or Asmanex. After 24 hours of in vitro release, CH-MF nanoparticles only

released 43% MF; in comparison, all MF from the Asmanex formulation was released

in 1 hour. In vitro release data show that chitosan coating allows much slower drug

release from nanoparticles. Drug release occurs by diffusion out of the nanoparticles as

well as due to degradation of chitosan and PLA. In the short term, diffusion processes









determine drug release, however over longer time periods degradation of chitosan and

PLA play a more dominant role. Similar observations were made from testosterone

contained in a PLGA film, initial release was as a result of diffusion from the film,

followed by a slower release due to the hydrolytic degradation of PLGA [101]. Both

chitosan coated and uncoated nanoparticles formulated with MF were able to release at

a slower rate compared to the formulation contained within the commercially available

Asmanex TwistnalerTM, currently used by oral inhalation for long-term asthma

management. PLGA microspheres of rifampicin coated with chitosan were shown to

produce a smaller burst effect compared to uncoated particles by Manca et al [97],

these results are in agreement with our observations from chitosan coated and

uncoated PLA nanoparticles.


Conclusion

Increased PLGA content resulted in a general increase in particles size with
significant increase in drug encapsulation efficiency and loading.

Increased PVA concentration resulted in reduced nanoparticle size,
encapsulation efficiency and drug loading.

Chitosan coated nanoparticles were not significantly larger than uncoated
particles, suggesting only a thin coating layer.

Reversal of zeta potential from negative for uncoated particles to positive for
chitosan coated particles confirmed presence of chitosan coating.

In vitro release studies showed that TA-PLGA nanoparticles released at a similar
rate to micronized drug particles.

In vitro release of MF-PLA and CH-MF produced a biphasic drug release profile.

In vitro release determined that CH-MF formulation released at the slowest rate,
with only 43% drug release over 24 hours.

Both chitosan coated (CH-MF) and uncoated MF-PLA nanoparticles released
drug at a slower rate compared to MF contained in the Asmanex TwisthalerTM.









* Chitosan coated particles, however, released at a much slower rate, only 20%
burst release was observed, followed by slow drug release.

* Chitosan coated and uncoated nanoparticles prepared with the corticosteroid MF
were able to release drug at a slower rate compared to MF contained in the
Asmanex TwisthalerTM.

* Chitosan coated MF-PLA nanoparticles were chosen for further development into
nanocomposite microspheres due to slow release rates observed.









Table 2-1. Advantages and
nanoparticles

Wet Milling



High pressure
homogenization


Supercritical fluid process












Solvent evaporation


disadvantages of various methods for production of


Advantages
Crystalline nanoparticles



40-500nm
Small or large scale


No organic solvent












Spherical nanoparticles
Monodispersed
nanoparticles
Ease of production
Suitable for hydrophobic
and hydrophilic drugs


Disadvantages
Contamination with grinding
material
Batch-to-batch variation
Microbial growth @300C
Aggregation and
coalescence of
nanosuspensions
Changes in crystallinity
Depends on efficiency of
atomization
Hydrophobic: hydrophilic
particles cannot be
produced inadequate
solvent systems
Poorly water soluble drugs
have poor CO2 solubility -4
low particle production
Resolved by high
temperatures
High pressures
Use of organic solvent
Not suitable for industrial
use














Sonicate at 60W for 5
mins



ONV microemulsion


Z


Evaporation of DCM


Centrifugation to
collect nanoparticles


4;


Lyophilization


Figure 2-1. Solvent evaporation technique to produce nanoparticles





























Figure 2-2. Schematic detailing methods to coat nanoparticles with chitosan








, >


Shake overnight
at room temp


Vacuum centrifuge


/


Add 1 ml
Mil :ie pliise
ACN:DDWX
60:40


S/-


R(.Tl mve
supernatant and
analyze by HPLC


S- 15

U~

'-sK-


Centrifuge to
remove
undissolved
polymer


Figure 2-3. Determine encapsulation efficiency and drug loading of nanoparticles











m 1% PVA

* 2% PVA

3% PVA


200


300
PLGA (mg)


400


Figure 2-4. Influence of PLGA and [PVA] on encapsulation efficiency of TA into
nanoparticles


100.0

90.0 -
90.0



70.0 -
80.0





50.0 -
70.0








30.0 -
60.0

50.0

40.0

30.0

20.0

10.0

0.0










250 -


200



- 150
N


S100 -
C-


50



0


m 3% PVA

S2% PVA

1% PVA


200


300


400


PLGA (mg)


Figure 2-5. Influence of PLGA and [PVA] on particle size of TA nanoparticles



































Figure 2-6. SEM of TA-PLGA formulated with 200mg PLGA and 1 % w/v PVA




































Figure 2-7. SEM of TA-PLGA formulated with 200mg PLGA and 2% w/v





































Figure 2-8. SEM of TA-PLGA formulated with 400mg PLGA and 1 % w/v PVA









100.00 -


80.00 -


60.00 -


5 40.00
micronizedd TA

20.00 300mg PLGA 2% PVA

300mg PLGA 1%PVA
0.00
0 100 200 300 400 500 600
Time (mins)

Figure 2-9. In vitro release TA-PLGA nanoparticles formulated with 300mg PLGA using
1% or 2% w/v PVA









120.00 -


100.00 -


80.00 -


"1 60.00


40.00 -
40.00 micronizedd TA

S400mg PLGA 1%PVA
20.00 -
A400mg PLGA 2%PVA

0.00 M
0 100 200 300 400 500 600
Time (mins)

Figure 2-10. In vitro release of TA-PLGA nanoparticles formulated with 400mg PLGA
and 1% or 2% w/v PVA




































Figure 2-11. SEM of MF-PLA nanoparticles incubated with chitosan 1% w/v









Table 2-2. Particle size distribution for chitosan coated MF-PLA nanoparticles prepared
with 0.1% or 1% w/v chitosan by either incubation or in situ coating with
chitosan
Method Particle size (nm) SD (nm)
Incubation 0.1% w/v chitosan Bimodal
Incubation 1% w/v chitosan Bimodal
In situ 0.1% w/v chitosan 226 10
In situ 1% w/v chitosan Bimodal









Table 2-3. Influence of chitosan coating to MF-PLA nanoparticles on
efficiency, drug loading, particle size and zeta potential
Drug Particle
loading SD size SD
Formulation (%w/w) (%w/w) (nm) (nm)
PLA nanoparticles 1.4 0.1 212 14


Chitosan coated PLA
nanoparticles


226


encapsulation

Zeta
potential
(mV)
-36.0

13.7







































Figure 2-12. SEM of uncoated MF-PLA nanoparticles formulated with 400mg PLA, 1%
w/v PVA







































Figure 2-13. SEM of chitosan coated MF-PLA nanoparticles formulated with 400mg
PLA, 1% w/v PVA and 0.1% w/v chitosan

















+{ '


+PLA-MF


CH-MF


m Asmanex


0 200 400


600


800


Time (mins)
Figure 2-14. In vitro release of MF-PLA, chitosan coated
micronized MF


1000


1200 1400


MF-PLA (CH-MF) and


120


100 -









CHAPTER 3
DEVELOPMENT OF NANOCOMPOSITE MICROSPHERES

Introduction

Although nanoparticles may be capable of providing increased pulmonary

retention and slow drug release, they are not efficiently deposited within the airways. In

order to deposit in the airways particles must have an MMAD in the respirable range of

1-5[pm [26]. Nanoparticles must first be incorporated into micron sized carrier particles,

upon exposure to the aqueous environment of the lung, these particles are released,

and deposit nanoparticles onto the lung epithelium [102]. Nanoparticles may be spray

dried with an excipient such as lactose to form nanocomposite microspheres with good

flow properties [103-105]. Ely et al developed effervescent spray dried nanoparticles,

these contain an active method to breakdown microspheres once deposited in the lung;

citric acid and carbonate react when delivered to the humid airways for form bubbles of

carbon dioxide and disperse nanoparticles [106]. Hadinoto et al investigated the

influence of nanoparticle size, chemical nature and feed concentration of polystyrene

and colloidal silica nanoparticles to be spray dried on the ability to form hollow

nanocomposite particles and found that a particular nanoparticle concentration

threshold must be reached in order to produce hollow microspheres composed of

nanoparticles alone [107]. Addition of phospholipids also reduced phagocytic uptake,

nanocomposite particles have been formulated by spray drying nanoparticles with

phospholipids, however these particles resulted in slower drug release after spray

drying [108].

Spray dried gelatin nanoparticles were found to be significantly larger following

spray drying under conventional conditions in a commercially available instrument,









however no size change was seen for polybutylcyanoacrylate particles [105, 109].

Spray dried PLGA dexamethasone nanoparticles were spray dried with 1,2-Dipalmitoyl-

sn-Glycero-3-Phosphocholine (DPPC) and hyaluronic acid. The resultant microspheres

produced released dexamethasone at a slower rate compared to nanoparticles alone

[103]. Aspirin nanoparticles spray dried with increasing concentrations of phospholipids

initially had similar burst release rates to nanoparticles alone, however, following the

initial burst nanocomposite particles with higher phospholipid contents released at a

slower rate [107]. Similar findings have also been noted in other studies [103]. Large

porous lactose particles have been developed by the Edwards group, the advantage of

these are reduced aggregation on storage compared to smaller non-porous particles

with an equivalent MMAD [102].

Lactose is a commonly used excipient for inhaled formulations, a number of

studies have used lactose to form nanocomposite microspheres [105, 106, 109]. PLGA

nanoparticles spray dried with trehalose or lactose are able to release nanoparticles on

contact with lung lining fluid. If these nanocomposite microspheres are produced by

spray drying using inlet temperatures of 100C or higher, it is no longer possible for

nanoparticles to be released from the formulation [110]. Commercially available spray

drying systems operate with inlet temperatures ranging from 100-220C, however, this

may be problematic for use with polymeric nanoparticles. PLGA and PLA which are

used to form nanoparticles have glass transition temperatures of 45-500C and 50-600C

respectively, thus spray drying at higher temperatures will result in changes in particle

shape and possibly agglomeration of nanoparticles [111].









Hypothesis

We hypothesize that a spray dryer capable of operating under lower temperatures

than commercially available will allow nanoparticles to be incorporated into

microspheres without changes in particle morphology, size or in vitro release

characteristics. During the drying of aerosol droplets within the spray dyer,

nanoparticles are protected from excessive heat due to evaporation of water. Once the

aerosol is dried, nanoparticles may be exposed to high temperatures resulting in

agglomeration of particles as well as changes in drug release rates, thus it is especially

important that the outlet temperature of the spray dryer should be lower than the Tg of

the polymer nanoparticles. Lactose will be used as an excipient; we hypothesize that

nanoparticles will be freely re-dispersed from nanocomposite microspheres composed

of lactose.

Materials and Methods

Chemicals

MF was donated by Ipca laboratories ltd (Mumbai, India). PLGA and PLA were

purchased from Lactel Absorbable Polymers (Pelham, AL, USA). Extra-fine lactose

was donated by EM industry (Hawthorne, NY, USA). DCM and ACN were purchased

from Fisher Scientific (Pittsburgh, PA, USA). PVA and medium molecular weight

chitosan were obtained from Sigma Chemical Co. (St. Louis, MO).

Development of Spray Dryer and Optimization of Operating Conditions

DL-PLA has a glass transition temperature (Tg) of 50-600C, a novel spray dryer capable

of operating at temperatures lower than the Tg of the polymers was developed (Figure

3-1); this allowed nanoparticles to be spray dried without causing particles to aggregate

during the process. Operating conditions of the spray dryer will be optimized to allow









dried microspheres to be collected using an Anderson cascade impactor (Copley

Scientific, UK). Initially, PLA-TA nanoparticles were spray dried with lactose to

determine the lowest temperatures and additional parameters under which dried

microspheres may be produced. Parameters such as sweep air flow, starting

temperature of the drying chamber, nebulizer air flow, nebulizer flow rate were varied.

Particle morphology was observed by SEM to determine particle morphology and

incorporation of nanoparticles into microspheres.

During method development, in order to determine operating parameters for spray

drying, total collection of TA-PLA nanoparticles on the ACI was evaluated by weight of

particles deposited on each stage. In addition, each stage was rinsed with 5 mL MeOH

and total TA content will be evaluated following dilution with DDW to a produce 50:50

MeOH:DDW solution. Total mass of nanoparticles collected on each stage was

determined. This allowed a fast screening method to determine the composition of

nanocomposite microspheres produced as well as an approximation of the MMAD of

the formulation.

The spray dryer design is shown in Figure 3-1. Warm dry air is directed into the heating

chamber from the bottom of the instrument. The port placed at the bottom of the

heating chamber is angled to allow the warmed dry sweep air to produce a vortex

through the heating chamber; preventing nebulized aerosol droplets impinging the spray

dryer walls. This results in reduced loss of formulation within the spray dryer, to

improve the final yield of spray dried nanoparticles. In addition to this, it increases the

time the aerosol droplets reside in the heating chamber, allowing a more gentle heat to

be applied. In addition width and height of the aerosol generated using the nebulizer









was measured, as a result the diameter of the drying chamber was designed to be

wider than the aerosol plume created by the glass nebulizer to reduce impaction and

loss to the chamber walls.

Inlet temperature was maintained in the range 75-800C while outlet temperature was

less than 350C. Parameters such as temperature, nebulizer air flow, flow rate of

nanoparticle suspension to be aerosolized and air flow through the spray drying device

were optimized in order to allow complete drying of the microspheres without subjecting

nanoparticles to temperatures above Tg. Optimal conditions were found to be with inlet

airflow of 15 L/min, 0.5 ml/min flow of nebulizer suspension with the whole system

under slightly negative pressure. Inlet temperatures less than 750C resulted in

incomplete drying of the microspheres, which could be observed by the collection of wet

particles in the ACI. Higher flow rates of the nebulized suspension also resulted in wet

particles being produced. The nanoparticle/lactose suspension is sprayed into a heated

drying chamber using a type A concentric circle glass nebulizer (Meinhard, Golden, CO)

at a rate of 0.5 mL/min under 20 psi with a corresponding air flow of 0.8 L/min, shown in

Figure 3-2. The nebulizer is able to operate at a maximal pressure of 30 psi, with a

linear correlation between air flow rate and pressure between 15 and 80 psi, (Air flow

(L/min) = 0.0328xPressure(psi) + 0.1401 with R2 = 0.9997). The heated chamber was

maintained under a slightly negative pressure in order to improve aerosol drying. Warm

dry sweep air at a flow rate of 15 L/min was used to create a vortex within the drying

chamber to prevent aerosol droplets from impinging onto the chamber walls. Inlet and

outlet temperatures were approximately 780C and <350C respectively; outlet

temperatures should be maintained below the Tg of the polymer in order to prevent









nanoparticle agglomeration. Typically commercial spray dryers operate at much higher

temperature, for example lactose is spray dried at an inlet temperature of 1600C and an

outlet temperature of 1050C, this method would be unsuitable for use with polymeric

nanoparticles [112]. An Anderson cascade impactor was set up in-line with the spray

dryer as a method of both collecting the spray dried formulation and also to selectively

collect the respirable fraction. Air flow through the impactor was set to 20 L/min,

following spray drying, the formulation was retrieved from stages 2-7. Particles

collected are expected to have a MMAD below 6.9pm.

Spray Dried Chitosan Coated MF-PLA Nanoparticles

Optimal spray dryer conditions determined were used for further development of

spray dried nanocomposite microspheres. In Chapter 2, chitosan coated PLA

nanoparticles were shown to produce the slowest drug release compared to uncoated

nanoparticles or MF contained in the Asmanex formulation (Figure 2-14), as a result

were spray dried with lactose. Chitosan coated nanoparticles (CH-MF) were suspended

in a lactose solution, then spray dried. Total solid feed content to be nebulized was

initially set to 5% w/v while the CH-MF: lactose composition was varied (10:90, 25:75,

50:50, 75:25 and 90:10). MMAD, particle morphology, nanoparticle entrapment

efficiency and in vitro release characteristics of spray dried CH-MF nanoparticles were

determined as described below.

Nanoparticle Entrapment Efficiency and Loading into Nanocomposite
Microspheres

In order to determine the percentage nanoparticle content of the nanocomposite

microspheres as well as the nanoparticle encapsulation efficiency 5 mg of the

nanocomposite microspheres were weighed and dissolved in DCM on the orbital shaker









for overnight. DCM was placed in the vacuum centrifuge (Jouan RC10.10) to dryness,

the dried residue was dissolved in 1 mL ACN:DDW 60:40 and analyzed by HPLC

(Hewlett Packard Series 1050). Drug loading and encapsulation efficiency were

determined by the equations below.

Actual MF (ug)
Drug loading (% w/w) = Ata----x 100
Weight of Formulation (ug)

Actual drug loading (% w/w)
Encapsulation Efficiency = ical drug loading (%x 100
Theoretical drug loading (%w/w)

Morphology of Nanocomposite Microparticles

SEM was performed to observe particle morphology using Scanning Electron

Microscope (SEM) JEOL JSM-6335F instrument (Major Analytical Instrument Center

(MAIC), UF, Gainesville, FL). The spray dried formulation was placed on a carbon stub

which was coated with carbon using a vacuum evaporator, SEM was performed at 2 kV.

Lactose alone was spray dried under the same conditions to allow comparison of

particle morphology of spray dried particles which do not contain any nanoparticles.

Determine M MAD of Nanocomposite Microparticles

Particle MMAD of spray dried hybrid particles was determined using an 8-stage, non-

viable Anderson cascade impactor (Copley Scientific), stainless steel collection plates

were used, the inhaler device was coupled to the impactor using a tailor made adapter.

The operational airflow used was 39 L/min for 6 seconds per actuation. Surfaces of the

particle collection sites on the ACI were coated with an ethanolic solution of brij 35 in

glycerol to avoid bias caused by particle bounce [113]. Each of the plates was coated

with 50 [L brij 35 solution, with the pre-seperator being coated with 100 [L to reduce

the bounce effect, ethanol will evaporate before HPLC analysis. The sample induction









port (SIP), high top, pre-separator and plates from each stage were washed with 5 mL

acetic acid buffer pH 4, to remove the chitosan coating, this was followed by incubation

with 5 mL acetonitrile to release MF from within the PLA matrix. The samples were

centrifuged then analyzed by HPLC as described above. Two capsules were half filled

with 10 mg of the spray dried formulation; the capsule was placed in the center chamber

of the device with the overlapping end pointing upwards and delivered via the

HandiHaler@ (Figure 3-4). Capsules were pieced in the HandiHaler, however the size

of the holes produced was designed to be smaller diameter than that used in the

commercial device, in order to prolong the residence/delivery time of the formulation

within the capsules and thereby allowing for a longer de-agglomeration process of the

formulation particles. The respirable fraction of the spray dried nanoparticles

formulation was compared with that of the commercially available MF formulation

delivered from the Asmanex inhaler device (Figure 3-5). The impactor will not give a

continuous distribution of particles, but will categorize particle size into certain size

ranges [114]. The respirable dose was calculated based on particles with MMAD <4.9

pm by ACI analysis, based on MF collection from stage 2 and below, based on the

particle size cutoff shown in Table 3-1. The Metered Dose (MD) was based on the MF

content using the actual amount of formulation weighed into the capsules for inhalation.

The Delivered Dose (DD) was determined based on the total recovered MF from all

stages of the ACI, pre-separator, Sample Induction Port (SIP), high top and adapter.

Drug remaining in the device and capsules were not included within the DD. Fine

particle dose (FPD) was determined by the summation of all MF collected from stage 2

and below and was compared to DD to determine FPF.









Fine Particle Dose (ug)
Fine Particle Fraction = x 100
Delivered Dose (ug)

In vitro drug Release Study

In vitro release was performed in 100 mL 1% v/v tween 80 in PBS, shaken at 30 rpm in

a hot shaker (Bellco biotechnology, Vineland, NJ) at 370C over 24 hours under constant

sink conditions, as previously described in Chapter 2. Tween 80 increased the

saturation concentration of MF in the dissolution media to 25 pg/ml at 370C. In order to

maintain sink conditions maximal concentrations obtained did not exceed 10% of the

saturation concentration. At specific time points (0, 15, 30, 90, 120, 180, 240, 360, 600

and 1440 minutes), 1 mL was removed using a pipette and subsequently filtered using a

0.02 |tm filter (Whatman Anotop plus filter), 1 mL fresh buffer was replaced into the

dissolution media. HPLC was used to quantify released drug concentrations using a

calibration curve ranging from 0.5-10 ug/mL with an R2>0.996. Spray dried CH-MF

formulations were compared to MF from the Asmanex TwisthalerTM.

Results and Discussion

Development of Spray Dryer and Optimization of Operating Conditions

A spray dryer capable of operating with inlet temperatures less than the Tg of

PLGA or PLA was designed. Additionally it was able to spray dry formulations using mg

quantities of materials. Initially PLA-TA nanoparticles were used to optimize the spray

dryer operational parameters, SEM of this formulation can be seen in Figure 3-9.

Figure 3-6 shows MMAD of particles collected directly from the ACI, a large fraction of

the total spray dried microspheres obtained are less than 5.8 pm, Figure 3-7 also shows

this as the total mass collected at each stage. Figure 3-8 shows that in general the total

nanoparticle: lactose content did not vary between the collection stages, with an









exception of a much lower incorporation of nanoparticles into microspheres <0.65 pm.

This may be due to exclusion of nanoparticles from microspheres that are very small, as

the nanoparticles are also sized at least 200 nm. Initially mass of formulation deposited

on the collection surfaces were determined by gravimetric analysis as a convenient

method to quantify larger amounts of collected formulation. However, for smaller

masses, it becomes more accurate to perform analysis by HPLC as gravimetric analysis

is not able to differentiate between formulation weight and water weight, therefore spray

dried CH-MF nanoparticles were further analyzed by HPLC [115].

Spray Dried Chitosan Coated MF-PLA Nanoparticles Influence of CH-MF:Iactose
Ratio on Morphology of Spray Dried Microspheres

Chitosan coated MF-PLA (CH-MF) nanoparticles were used to optimize

composition of nanocomposite microspheres. The effect of increasing

nanoparticle:lactose ratio on particle morphology was investigated. CH-MF nanoparticle

content was varied at 10%, 25%, 50% and 75% of the total feed of 5% w/v being spray

dried. Higher nanoparticle content was not possible as a result of blockage of the

nebulizer nozzle with the use of 90% nanoparticle composition. At low composition of

10% nanoparticles, microspheres formed were spherical as shown in Figure 3-10,

however, as nanoparticle content increased, particles appeared more deflated and

hollow as shown in Figures 3-11, 3-12 and 3-13. Advantages of spray dried hybrid

particles containing higher nanoparticle content include a lower mass to be inhaled as

well as production of hollow particles which in turn will have lower MMAD as in

comparison to solid particles with similar particle size. Hollow particles tend to have

larger particle sizes compared to non-porous particles of the same MMAD. This

property allows deposition of these particles in the same regions of the airways,









however less aggregation should be expected on storage. As a result of this, chitosan

coated nanoparticles were developed using a high nanoparticle: lactose ratio of 75:25.

The high solid feed content of 5% w/v tended to lead to clogging of the nebulizer nozzle,

therefore particles were spray dried with a solid feed concentration of 1.25% w/v.

Lactose alone was spray dried at a concentration of 1.25% w/v to determine surface

morphology of these microspheres in the absence of incorporated nanoparticles. Figure

3-14 shows the SEM of spray dried lactose alone, microspheres have a smooth surface

in comparison to lactose spray dried with nanoparticles. The resultant microspheres

were still produced using a ratio 75% CH-MF nanoparticles and 25% lactose, therefore

the final formulation will have the same nanoparticle and drug loading.

Spray Dried Chitosan Coated MF-PLA Nanoparticles Nanoparticle Incorporation

Chitosan coated nanoparticles were spray dried with lactose 75% and 25%

respectively, with a solid feed concentration of 1.25% w/v. Solid feed content was

reduced from 5% w/v from previous studies with uncoated MF-PLA nanoparticles to

1.25% w/v as chitosan coated nanoparticles tended to clog the nebulizer at high

concentrations. A relatively high nanoparticle: lactose ratio was chosen in order to

produce a formulation which contained a clinically appropriate dose and could be

administered by one or two inhalations after filling capsules for use in the HandiHaler@.

Theoretical MF loading into the spray dried CH-MF formulation is 1.5% w/w, actual MF

loading was very similar at 1.43% w/w (SD=0.08% w/w). It was determined that 96%

(SD=6%) of the total nanoparticles used were incorporated into the final spray dried

formulation. SEM of the formulation contained within the Asmanex TwisthalerTM is

shown in Figure 3-15, this shows a blend of lactose with MF.









Spray Dried Chitosan Coated MF-PLA Nanoparticles MMAD

MMAD of spray dried CH-MF nanoparticles were characterized by ACI to

determine FPF of the delivered dose from the HandiHaler at 39 L/min for 6 seconds.

The United States Pharmacopeia recommends that DPIs should be tested at a flow rate

which corresponds to a pressure drop across the inhaler of 4 kPa [116]. As the

HandiHaler@ produces a pressure drop of 4 kPa at 39 L/min, this was the air flow

tested. On inspiration the capsule vibrates inside the center chamber, this results in

mechanical agitation of powder aggregates contained and results in their break up

[117]. Increasing retention of the formulation within the capsule before delivery to the

lung allows increased time over which de-aggregation may occur. In order to increase

the time that the spray dried formulation is mechanically agitated within the device;

holes used to pierce the capsule were customized to be smaller than those usually

made with the HandiHaler, while allowing the capsule to empty with one inhalation.

FPF of MF contained in Asmanex TwisthalerTM was used as a control for the spray

dried CH-MF nanoparticles. The Metered Dose (MD) and Delivered Dose (DD) of MF

delivered from the Asmanex TwisthalerTM produced by Schering-Plough is available as

220 pg or 440 pg and 200 pg or 400 pg respectively [118].

The HandiHaler is a high efficiency inhaler and was used as it is capable of

delivering large doses up to 50 mg by inhalation [119]. This allows a clinical dose of MF

from the spray dried chitosan coated nanoparticle formulation to be administered with

one or two inhalations. In addition, prepared spray dried formulations can be packaged

into capsules with relative ease in comparison to other DPI devices such as the

FlixotideTM DiskhalerTM which would require special blister packing equipment to be

used [120].









FPF on DD from the spray dried chitosan coated nanoparticle formulation was 14.6%

(SD=4.3%), whereas the Asmanex TwisthalerTM had 20.0% (SD=4.7%). Three

batches of spray dried formulation were compared in order to determine batch to batch

variability of the spray drying process (n=3 per batch), shown in Figure 3-16.

MMAD of MF from the TwisthalerTM was evaluated in order to make comparisons with

the FPF on DD of the spray dried nanoparticle formulation with a known commercial

formulation. Literature values for FPF of the Asmanex TwisthalerTM are reported to

be 39.9% (SD=2.5) and 35.6% (SD=3.4) at 200 pg and 400 pg DD with an airflow of 60

L/min. The fine particle dose was however defined as particle with MMAD <6.5 pm

[121]. Differences between ACI conditions will account for the lower FPF observed with

the Asmanex compared to the literature reported values. To summarize, FPF of the

DD from the spray dried CH-MF formulation is similar to commercially available MF from

Schering-Plough's Asmanex TwisthalerTM.

Spray Dried Chitosan Coated M F-PLA Nanoparticles In vitro release

MF blended with lactose was removed from the Asmanex TwisthalerTM, Figure 3-

15. At 1 hour 100% of MF from the Asmanex is released, in comparison, spray dried

CH-MF exhibits an initial burst, followed by slower release. Spray dried CH-MF shows

a slightly greater burst effect compared to CH-MF nanoparticles alone of 30% and 20%

respectively. At 24 hours approximately 70% of the total MF contained within the spray

dried formulation is released, however, CH-MF nanoparticles alone only released 43%

of the total MF content at this time point (Figure 3-17). Lyophilized nanoparticles must

be re-dispersed in a lactose solution before they can be spray dried. It is possible that

during this process, MF is able to disperse to the surface of the CH-MF nanoparticles,

leading to a greater burst release. Following the initial phase, release rate from CH-MF









nanoparticles and spray dried formulation were similar, indicating that nanoparticles did

not aggregate during the spray drying process. In vitro release of CH-MF nanoparticles

did not change following spray drying. Additionally, the spray dried formulation released

MF at a much slower rate compared to the commercial MF prepared in the Asmanex

TwisthalerTM. Studies conducted by other groups observed slower drug release from

PLGA nanoparticles spray dried with lipids, however it is unclear if this is a result of high

temperatures used or a property of the lipids used [102, 107].

Conclusion

The spray dryer was designed to operate at outlet temperatures below the Tg of
PLA or PLGA.

Spray dried particle morphology was influenced by nanoparticle: lactose
composition used.

Spherical microspheres with MMAD in the respirable range were produced
containing nanoparticles.

Increase in nanoparticle: lactose ratio resulted in changes in microsphere
morphology.

Lower nanoparticle:lactose ratios produced spherical microspheres whereas
higher nanoparticle concentrations produced hollow particles as shown by SEM.

Spray dried chitosan coated MF-PLA nanoparticles composed of 75%
nanoparticles and 25% lactose are able to provide a clinically appropriate dose in
approximately two inhalations using 20 mg of the formulation from the
HandiHaler.

FPF on DD of the spray dried chitosan coated MF-PLA nanoparticles was 14.6%
(SD=4.3%), whereas the Asmanex TwisthalerTM was 20.0% (SD=4.7%).

In vitro release of MF from the spray dried chitosan coated MF-PLA
nanoparticles was much slower than from the formulation contained within the
Asmanex TwisthalerTM.

A biphasic release profile was observed for spray dried CH-MF hybrid particles









* After CH-MF nanoparticles were spray dried, initial burst release observed was
faster than for CH-MF nanoparticles alone. Following the initial burst however,
the release rates were found to be similar.








































Figure 3-1. Spray dryer, designed to operate under lower temperatures than
commercially available instruments
















Flush Capllk
Lapped N=2


Figure 3-2. TR-30-A3 nebulizer, type A flush capillary lapped nozzle 30 psi, 3 mL/min
[122]































Figure 3-3. Schematic to determine incorporation efficiency of nanoparticles into spray
dried formulation

















Figure 3-4. HandiHaler (1) dust cap, (2) mouthpiece, (3) mouthpiece ridge, (4) base,
(5) green piercing button, (6) center chamber, (7) air intake vents [123]











WI
C


Mouthpiece


Fluted Chimney o
Nozzle
Swirl Chamber J



Reservoir

Dose Plate






. Base


Figure 3-5. Cross section of Asmanex TwisthalerTM [121]









3-1. MMAD


cutoff for ACI


analysis based on air flow rate


L/min 20.0 22.5 28.3 30.0 39.0 45.4 60.0
Stage


17.8-
10.7
10.7-6.9
6.9-5.6
5.6-3.9
3.9-2.5
2.5-1.3
1.3-0.8
0.8-0.5
<0.5


16.8-
10.1
10.1-6.5
6.5-5.3
5.3-3.7
3.7-2.4
2.4-1.2
1.2-0.8
0.8-0.4
<0.4


15.0-9.0 14.6-8.7 12.8-7.7 11.8-7.1 10.3-6


9.0-5.8
5.8-4.7
4.7-3.3
3.3-2.1
2.1-1.1
1.1-0.7
0.7-0.4
<0.4


8.7-5.6
5.6-4.6
4.6-3.2
3.2-2.0
2.0-1.1
1.1-0.7
0.7-0.4
<0.4


7.7-4.9
4.9-4.0
4.0-2.8
2.8-1.8
1.8-0.9
0.9-0.6
0.6-0.3
<0.3


7.1-4.6
4.6-3.7
3.7-2.6
2.6-1.7
1.7-0.9
0.9-0.6
0.6-0.3
<0.3


6.2-4.0
4.0-3.0
3.0-2.2
2.3-1.0
1.4-0
0.8-0
0.5-0
<0.3


Table


(L/min)


1
2
3
4
5
6
7
Filter










30 -

S25 -

. 20 -

F5 15 -

E 10 -
0
o 5
" 0


Figure 3-6. MMAD of nanocomposite PLA-TA particles


I I I I I
5.8-9 4.7-5.8 3.3-4.7 2.1-3.3 1.1-2.1 0.65-1.1 0.43-0.65

M MAD (um)










12 -


10
Cm
E
8

E
a 6
0

E
0
u- 2

0
5.8-9 4.7-5.8 3.3-4.7 2.1-3.3 1.1-2.1 0.65-1.1 0.43-0.65
MMAD (pm)

Figure 3-7. Total PLA-TA nanoparticles for each MMAD range









10.00 -
9.00 -
8.00 -
00
a 7.00 -
.2'g 6.00

5.00

.E 4.00 -
0U 3.00 -
m O
M m 2.00 -

1.00 -
0.00
5.8-9 4.7-5.8 3.3-4.7 2.1-3.3 1.1-2.1 0.65-1.1 0.43-0.65
MMAD (um)


Figure 3-8. Percentage of PLA-TA nanoparticles at each particle cutoff compared with
total weight







































Figure 3-9. SEM of PLA-TA nanoparticles spray dried with lactose






































Figure 3-10. Spray dried chitosan coated MF-PLA nanoparticles containing 10% of total
solid feed as nanoparticles






































Figure 3-11. Spray dried chitosan coated MF-PLA nanoparticles containing 25% of total
solid feed as nanoparticles





































Figure 3-12. Spray dried chitosan coated MF-PLA nanoparticles containing 50% of total
solid feed as nanoparticles






































Figure 3-13. Spray dried chitosan coated MF-PLA nanoparticles containing 75% of total
solid feed as nanoparticles







































Figure 3-14. Spray dried lactose, 1.25% w/v solid feed content


100








































Figure 3-15. SEM of formulation contained in Asmanex TwisthalerTM









30 -
28 -
26 -
24 -
.22 -
20 -
S18 -
0 16 -
14 -
S12 -
S 10 -
8
6
4
2
0
Batch 1 Batch 2 Batch 3




Figure 3-16. FPF of DD (%) of spray dried chitosan coated nanoparticles, to compare
batch to batch variability (n=3)


102












100l OO


*Spray dried CH-MF nanoparticles
EAsmanex MF


CH-MF


0 200 400


600


800


1000


1200


1400


Time (mins)


Figure 3-17. In vitro release of MF from spray dried CH-MF nanoparticles and
Asmanex


103









CHAPTER 4
SUMMARY

Polymeric nanoparticles containing PLGA and the model drug TA were developed

using the solvent evaporation. Spherical unimodal TA-PLGA nanoparticles were

produced using this technique. It was shown that various process parameters could

influence particle size, drug encapsulation and drug release. Increasing polymer

content was shown to increase encapsulation of TA in TA-PLGA nanoparticles and

result in a trend of increasing particle size. As PVA surfactant concentration was

increased, encapsulation efficiency decreased and smaller nanoparticles were formed.

These TA-PLGA nanoparticles were able to release at a similar rate as micronized TA.

Further developments were made using PLA to form nanoparticles. Mometasone

furoate was used as this is a more lipophilic drug and would show slower drug release

from polymer nanoparticles. Drug release was further reduced by coating MF-PLA

nanoparticles with chitosan. MF released at a substantially slower rate from chitosan

coated nanoparticles compared with MF-PLA nanoparticles and micronized MF that was

used as a control.

Nanocomposite particles were developed using a spray dryer that was developed;

this spray dryer operated with outlet temperatures below the Tg of PLA or PLGA making

it suitable for spray drying polymer nanoparticles. The spray dryer was able to produce

particles with MMAD in the respirable range. Nanocomposite microspheres were able

to deliver a clinically relevant dose with two SPIRIVA capsules filled with 10 mg of the

formulation using the HandiHaler device. SPIRIVA capsules were placed in the

center chamber of the HandiHaler and pierced by pressing the button to a pre-

determined level. Holes size produced in the capsule was customized to increase


104









retention within the device for a longer time period and therefore increase the time over

which de-aggregation may occur. As a result of both the formulation and optimizing the

inhaler device for use with this formulation, the FPF was similar to that from the

Asmanex TwisthalerTM.

MF release from the spray dried CH-MF was compared to the nanoparticles before

spray drying and also MF formulation contained within the Asmanex TwisthalerTM. MF

was release rapidly from the Asmanex with 100% release by 1 hour. In comparison

both CH-MF and the spray dried formulation exhibited a biphasic release profile. CH-

MF nanoparticles showed a burst release of approximately 20% compared to 30% from

the spray dried formulation. A possible explanation for this may be as a result of having

to re-disperse lyophilized nanoparticles in a lactose solution before spray drying. This

could cause some of the MF to diffuse to closer to the surface of the nanoparticles and

result in a greater burst effect observed with spray dried nanoparticles. Following the

initial burst, both the CH-MF nanoparticles and spray dried formulation release with a

similar rate. In vitro release from nanoparticles was not altered by the spray drying

process.


105









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BIOGRAPHICAL SKETCH

Gina Patel was born in 1981, London, England. She graduated from the

University of Bath, Bath, UK in 2003 as a Master of Pharmacy. During her studies she

participated in an internship at the University of Florida under the supervision of Dr

Hochhaus. After qualification as a pharmacist she worked as a hospital pharmacist at

Musgrove Park Hospital, Taunton, UK before joining the Department of Pharmaceutics,

University of Florida.





PAGE 1

1 DEVELOPMENT OF NANOCOMPOSITE CORTICOSTEROID PARTICLES FOR USE IN ASTHMA By GINA PATEL A DISSERTATION PRESENTED TO THE GRADUATE SCHOOL OF THE UNIVERSITY OF FLORIDA IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF DOCTOR OF PHILOSOPHY UNIVERSITY OF FLORIDA August 2010

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2 20 10 Gina Patel

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3 This work is dedicated to my Mum, grandparents, husband and family, for their love and support.

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4 ACKNOWLEDGMENTS I would like to express my sincere gratitude and appreciation to my mentor Dr G nther Hochhaus, for accepting me into his group and his guidance and support during the time I worked with him. I would also like to thank my committee members, Dr Hartmut Derendorf, Dr Anthony Palmieri III, Dr Jeffrey Hughes and Dr Christopher Batich for their guidance on this project. I thank Dr Arun Ranade for all the invaluable advice given to me in guiding the development of the spray dryer device. During my studies I have been fo rtunate enough to work as a teaching assistant for Dr Cary Mobley, I would like to thank him for making this time enjoyable and for always having the time to help and advise on my research. I would like to thank everyon e in the D epartment of Pharmaceutics I would like to thank Yufei Tang for all the invaluable advice she gave me. I am grateful to all of the past and present members of the Hochhaus group, Isabel Andueza, Vikram Arya, Intira Coowantiw ong, Navin Goyal, Manish Issar, Bhargava Kandala, Keerti Mudunuri, Elanor Pinto, Srikumar Sahasranaman, Wan Sun, Nasha Wang, Benjamin Webber, Yanning Wang and Kai Wu. For help with the experimental work over the years, I would like to thank all of my form er assistants, Christian Diestelhorst, Anica Liero, Gesa Nippel and Pooja Patel I must also thank Marc Rohrschneider for his advice and help on many aspects of this project I th ank them all for their friendship, caring and support over the years for w hich I am truly grateful. I would also like to show my gratitude and appreciation to Doug and Diane Ried for welcoming me into their family while I participated in an internship in the College of Pharmacy.

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5 I would like to thank my M u m, grandparents, husban d Jason Kwan and family for all their love and support throughout my education. Without their encouragement, none of this would have been possible.

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6 TABLE OF CONTENTS page ACKNOWLEDGMENTS ...................................................................................................... 4 LIST OF TABLES ................................................................................................................ 8 LIST OF FIGURES .............................................................................................................. 9 LIST OF ABBREVIATIONS .............................................................................................. 11 ABSTRACT ........................................................................................................................ 14 C H APT ER 1 INTRODUCTION ........................................................................................................ 17 Asthma ........................................................................................................................ 17 Asthma Treatment ...................................................................................................... 17 Adverse Effects of ICS ............................................................................................... 17 Fate of ICS .................................................................................................................. 18 Pulmonary Targeting of ICS ....................................................................................... 19 Ideal Corticosteroid ..................................................................................................... 20 Inhaler Device ............................................................................................................. 21 Influence of Mass Median Aerodynamic Diameter .................................................... 23 Control led Release Inhaled Formulations .................................................................. 23 Polymeric Nanoparticles ...................................................................................... 25 Solid Lipid Nanoparticles (SLN) and Microparticles ........................................... 26 Low Density Microspheres ................................................................................... 26 Pulsed Laser Deposition (PLD) ........................................................................... 27 Oligolactic Acid (OLA) .......................................................................................... 28 Pro -drugs .............................................................................................................. 28 Ester Formation .................................................................................................... 29 Liposomes ............................................................................................................ 30 Enhancing Mucoadhesive Properties .................................................................. 31 Objectives ................................................................................................................... 31 2 DEVELOPMENT AND CHARACTERIZATION OF SLOW RELEASE POLYMERIC CORTICOSTEROID NANOPARTICLES ............................................ 39 Introduction ................................................................................................................. 39 Hypothesis .................................................................................................................. 42 Materials and Methods ............................................................................................... 43 Chemicals ............................................................................................................. 43 Nanoparticle Preparation ..................................................................................... 43 Drug loading and Encapsulation Efficiency ......................................................... 45 Particle Size, Zeta Potential and Morphology ..................................................... 46

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7 In vitro Drug Release Study ................................................................................. 46 Results and Discussion .............................................................................................. 47 Influence of PLGA and PVA on Encapsulation Efficiency and Particle Size of Nanoparticles ................................................................................................ 47 Influence of PLGA and PVA on Morphology of Nanoparticles ........................... 48 In Vitro Release Study of TA -PLGA Nanoparticles ............................................ 49 Chito san Coated and Uncoated PLA MF Nanoparticles .................................... 50 Conclusion .................................................................................................................. 52 3 DEVELOPMENT OF NANOCOMPOSITE MICROSPHERES ................................. 71 Introduction ................................................................................................................. 71 Hypothesis .................................................................................................................. 73 Materials and Methods ............................................................................................... 73 Chemicals ............................................................................................................. 73 Development of Spray Dryer and Optimization of Operating Conditions .......... 73 Spray Dried Chitosan Coated MF -PLA Nanoparticles ........................................ 76 Nanoparticle Entrapment Efficiency and Loading into Nanocomposite Microspheres .................................................................................................... 76 Morphology of Nanocomposite Microparticles .................................................... 77 Determine MMAD of Nanocomposite Microparticles .......................................... 77 In vitro drug Release Study ................................................................................. 79 Results and Discussion .............................................................................................. 79 Development of Spray Dryer and Optimization of Operating Conditions .......... 79 Spray Dried Chitosan Coated MF -PLA Nanoparticles Influence of CH MF:lactose Ratio on Morphology of Spray Dried Microspheres ..................... 80 Spray Dried Chitosan Coated MF -PLA Nanoparticles Nanoparticle Incorporation ..................................................................................................... 81 Spray Dried Chitosan Coated MF -PLA Nanoparticles MMAD ......................... 82 Spray Dried Chitosan Coated MF -PLA Nan oparticles In vitro release ........... 83 Conclusion .................................................................................................................. 84 4 SUMMARY ................................................................................................................ 104 LIST OF REFERENCES ................................................................................................. 106 BIOGRAPHICAL SKETCH .............................................................................................. 117

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8 LIST OF TABLES Table page 1 -1 PKPD properties of inhaled corticosteroids ........................................................... 34 1 -2 Average AU Cs (n=3) in the lung, liver and brain and pulmonary targeting (PT) in neonatal rats after intratracheal administration (50 g/kg) of uncoated budesonide and PLA coated budesonide.............................................................. 36 1 -3 Cumulative Receptor Occupancy (AUC), Pulmonary Targeting and Mean Pulmonary Effect Times (MET) After Intratracheal Administration of Escalating Doses of TAP in 800 nm Liposomes ................................................... 37 1 -4 Influence of liposome composition on mucoadhesion and zeta potential ............ 38 2 -1 Advantages and disadvantages of various methods for production of nanoparticles .......................................................................................................... 54 2 -2 Particle size distribution for chitosan coated MF PLA nanoparticles prepared with 0.1% or 1% w/v chitosan by either incubation or in situ coating with chitosan .................................................................................................................. 66 2 -3 Influence of chitosan coating to MF -PLA nanoparticles on encapsulation efficiency, drug loading, particle size and zeta potential ...................................... 67 3 -1 MMAD cutoff for ACI analysis based on air flow rate (L/min) ............................... 91

PAGE 9

9 LIST OF FIGURES Figure page 1 -1 Fate of Inhaled Corticosteroids .............................................................................. 33 1 -2 Pulsed laser deposition (PLD) Nanoclusters of polymer from the target are deposited on larger micronized drug particles as continuous coatings that sustain the release rate of drug in solution ........................................................... 35 2 -1 Solvent evaporation technique to produce nanoparticles ..................................... 55 2 -2 Schematic detailing methods to coat nanoparticles with chitosan ....................... 56 2 -3 Determine encapsulation efficiency and drug loading of nanoparticles ............... 57 2 -4 Influence of PLGA and [PVA] on encapsulation efficiency of TA into nanoparticles .......................................................................................................... 58 2 -5 Influence of PLGA and [PVA] on particle size of TA nanoparticles ...................... 59 2 -6 SEM of TA-PLGA formulated with 200mg PLGA and 1% w/v PVA ..................... 60 2 -7 SEM of TA-PLGA formulated with 200mg PLGA and 2% w/v .............................. 61 2 -8 SEM of TA-PLGA formulated with 400mg PLGA and 1% w/v PVA ..................... 62 2 -9 In vitro release TA -PLGA nanoparticles formulated with 300mg PLGA using 1% or 2% w/v PVA ................................................................................................. 63 2 -10 In vitro release of TA -PLGA nanoparticles formulated with 400mg PLGA and 1% or 2% w/v PVA ................................................................................................. 64 2 -11 SEM of MF -PLA nanoparticles incubated with chitosan 1% w/v .......................... 65 2 -12 SEM of uncoated MF -PLA nanoparticles formulated with 400mg PLA, 1% w/v PVA ......................................................................................................................... 68 2 -13 SEM of chitosan coated MF PLA nanoparticles formulated with 400mg PLA, 1% w/v PVA and 0.1% w/v chitosan ...................................................................... 69 2 -14 In vitro release of MF -PLA, chitosan coated MF -PLA (CH MF) and micronized MF ........................................................................................................ 70 3 -1 Spray dryer, designed to operate under lower temperatures than commercially available instruments ....................................................................... 86 3 -3 Schematic to determine incorporation efficiency of nanoparticles into spray dried formulation ..................................................................................................... 88

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10 3 -4 HandiHaler ........................................................................................................... 89 3 -5 Cross section of Asmanex TwisthalerTM ............................................................. 90 3 -6 MMAD of nanocomposite PLA -TA particles .......................................................... 92 3 -7 Total PLA -TA nanoparticles for each MMAD range .............................................. 93 3 -8 Percentage of PLA -TA nanoparticles at each particle cutoff compared with total weight ............................................................................................................. 94 3 -9 SEM of PLA-TA nanoparticles spray dried with lactose ....................................... 95 3 -10 Spray dried chitosan coated MF -PLA nanoparticles containing 10% of total solid feed a s nanoparticles .................................................................................... 96 3 -11 Spray dried chitosan coated MF -PLA nanoparticles containing 25% of total solid feed as nanoparticles .................................................................................... 97 3 -12 Spray dried chitosan coated MF -PLA nanoparticles containing 50% of total solid feed as nanoparticles .................................................................................... 98 3 -13 Spray dried chitosan coated MF -PLA nanoparticles containing 75% of total solid feed as nanoparticles .................................................................................... 99 3 -14 Spray dried lactose, 1.25% w/v solid feed content ............................................. 100 3 -15 SEM of formulation contained in Asmanex TwisthalerTM ................................. 101 3 -16 FPF of DD (%) of spray dried chitosan coated nanoparticles, to compare batch to batch variability (n=3) ............................................................................. 102 3 -17 In vitro release of MF from spray dried CH -MF n anoparticles and Asmanex 103

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11 LIST OF ABBREVIATION S ACI Anderson cascade impactor ACN Acetonitrile AUC Area under the curve AUCbrain Area under the receptor occupancy time profile of the brain AUCliver under the re ceptor occupancy time profile of the liver AUClung under the receptor occupancy time profile of the lung BDP Beclomethasone dipropionate BMP Beclomethasone monopropionate BUD Budesonide CFC Chlorofluorcarbon CH-MF Chitosan coated mometasone furoate poly( lactic acid ) nanoparticles CIC Ciclesonide CIC -AM Ciclesonide active metabolite CL Clearance COPD Chronic obstructive pulmonary disease DCM Dichloromethane DDW Double distilled water DD Delivered dose Des -CIC Des isobutyryl ciclesonide DPI Dry powder inhaler DPPC 1,2-Dipalmitoyl -sn -Glycero -3 Phosphocholine EE Encapsulation efficiency F Bioavailability FDA Federal drug administration

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12 FLU Flutamide Foral Oral bioavailability FP Fluticasone propionate FPF Fine particle fraction fu Fraction of drug unbound g Gram GINA Global initiative for asthma HFA Hydrofluroalkane HPLC High performance liquid chromatography (Ultraviolet detection) ICS Inhaled corticosteroid kV Kilo volt MAIC Major Analytical Instrument Center MD Metered dose MDI Metered dose inhaler MeOH Methanol MF Mometasone furoate MF -PLA Mometasone furoate and poly( lactic acid ) nanoparticles mL Milliliter ms Milisecond MW Molecular weight nm Nanometer ODS Octadecyl silane PBS Phosphate buffered saline PLA Poly (lactic acid ) PLD Pulsed laser deposition

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13 PLGA Poly (lactic -co glycolic acid ) pMDI Pressurized metered dose inhaler PT Pulmonary targeting PVA Poly vinyl alcohol RRA Relative receptor affinity RPM Rotations per minute SD Standard deviation SEM Scanning electron microscope SMI Soft Mist Inhaler TA Triamcinolone acetonide TAP Triamcinolone acetonide phosphate TA -PLGA T riamcinolone acetonide and poly( lactic -co -gly colic acid ) nanoparticles TAP-lip Triamcinolone acetonide phosphate liposomes TAP-sol Triamcinolone acetonide phosphate sol ution Tg Glass transition temperature Vdss Volume of distribution at steady state W Watt % v/v % volume in volume % w/v % weight in volume % w/w % weight in weight

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14 Abstract of Dissertation Presented to the Graduate School of the University of Florida in Partial Fulfillment of the Requirements for the Degree of Doctor of Philosophy DEVELOPMENT OF NANOCOMPOSITE CORTICOSTEROID PARTICLES FOR USE IN ASTHMA By Gina Patel August 2010 Chair: G nther Hochhaus Majo r: Pharmaceutical Sciences Asthma is a chronic inflammatory condition of the airways resulting in episodes wheezing and breathlessness. Inhaled Corticosteroids (ICS) are used to prevent the occurrence of asthma attacks. The clinical effect of ICSs depen ds on the time the drug resides in the lung therefore increasing the drug residence time in the lung improves asthma therapy. It has been proposed that nanoparticles could escape clearance mechanisms in the lung and adhere strongly to the lung surface, leading to increased residence time. There are two main barriers to this approach, first ly nanoparticles cannot deposit in the lung, and instead they are exhaled. Second ly the particles must be formulated to release drug slowly in order to take advantage of the increased residence time. In order to further improve lung targeting of corticosteroids, poly(lactic -co glycolic acid) PLGA and poly(lactic acid) PLA were used to produce polymeric nanoparticles of Triamcinolone Acetonide (TA) and Mometasone Furoate (MF) using the solvent evaporation technique. TA was used as a model corticosteroid to allow optimization of various production parameters to produce PLGA nanoparticles. The solvent evaporation technique was used to produce polymeric nanoparticle s. A number of

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15 process parameters may be varied to alter nanoparticle size and drug encapsulation efficiency. Polyvinyl alcohol ( PVA) s urfactant concentration and PLGA content were varied to determine their influence on particle size, drug encapsulation efficiency, particle morphology and in vitro release characteristics of TA nanoparticles Increasing PLGA content resulted in a trend of increasing particle size and dru g encapsulation. As PVA concentration was increased particles tended to reduce in size and drug loading. Nanoparticles produced ranged in particle size between 156-20 9 nm. In addition when a low concentration of 1 % or 2% w/v PVA was used to produce nanoparticles combined wit h the use of only 200 mg PLGA TA crystals were observed by scanning electron microscopy. In vitro release studies revealed TA -PLGA nanoparticles released drug at a similar rate to micronized TA with 50% drug release being observed within 15 minutes In order to produce nanoparticles which can deliver drug at a slower rate compared to micron sized particles, a number of changes can be made to nanoparticle production; a more lipophilic drug MF in combination with a more hydrophobic polymer, PLA can be used to further slow down drug release. Subsequently MF nanoparticles (MF PLA) were using 10 mg MF, 400 mg PLA and 1% w/v PVA these particles showed slow release compared to MF contained in the Asmanex TwisthalerTM. To further reduce MF release rate, nanoparticles were coated with chitosan. In vitro release studies showed that chitosan coated MF -PLA nanoparticles (CH MF) showed significantly slower release compared to both uncoated nanoparticles and MF contained within the Asmanex TwisthalerTM. In vitro release studies determined 100% MF occurred after 1 hour for the Asmanex formulation, in comparison at this time only 50% and 24% MF was released from MF -PLA and CH MF respectively.

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16 A novel spray dryer was designed so that operating conditions of t his device allowed outlet temperatures to remain below the glass transition temperature of PLGA and PLA, thus making this system suitable to spray dry polymeric nanoparticles. Incorporation of these nano particles into lactose based microspheres by spray drying resulted in spherical nanocomposite microspheres. Analysis of these microspheres showed complete incorporation of nanoparticles into the formulation. Optimal conditions for incorporation of nanoparticles into microspheres were using a composition o f 75 % nanoparticles and 25 % lactose. The fine particle fraction of the microspheres was comparable to that of MF from the Asmanex TwisthalerTM. A biphasic release of MF was observed from the microspheres, with a significantly slower release compared to Asmanex. The spray drying process did not seem to alter the release properties of chitosan coated nanoparticles.

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17 CHAPTER 1 INTRODUCTION Asthma In the United States (US) asthma affects 16.4 million adults and 7 million children and is increasing in prevalence [1 -3] Asthma can be defined as a complex inflammatory dis ease of the airways involving many inflammatory mediators. Asthma is a chronic reversible inflammatory disease of the airways that is characterized by episodes of wheezing and breathlessness [4] A component of asthma involves chronic inflammation, if left untreated this can progress to airway remodeling, result ing in irreversible airway narrowing [5] Asthma Treatment The Global Initiative for Asthma (GINA) was established in 1993 in order to focus on reducing asthma prevalence, morbidity and mortality throughout the world. Published guidelines for the treatment and prevention of asthma were produced by GINA Current therapy includes Inhaled Corticosteroids (ICS), longacting 2 agonists, leukotriene modifiers, theophylline and short acting 2 agonists for immediate relief [4] ICS therapy targets the underlying inflammation pres ent in asthma, to reduce long term consequences such as progression of asthma to reversible obstruction in the airways. ICS therapy is therefore the cornerstone for asthma treatment and is recommended for all severity levels of persistent asthma by GINA [4, 6] Adverse E ffects of ICS Pulmonary drug delivery has many attractive prospects for local and systemic action. Select ively targeting drug delivery to the lung allows lower doses to be

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18 administered providing a reduction in systemic side effects whilst maintaining clinical efficacy Although ICS are considered the gold standard in asthma management they are responsible for many local and systemic adverse effects. Local side effects include oropharyngeal candidiasis, dysphonia, perioral dermatitis (which may occur with the use of spacer devices or nebulizers attached to face masks ), cough, thirst and in r are cases tongue hypertrophy. Additionally s ystemic side effects include adrenal insufficiency, cataracts, glaucoma, growth retardation in children, osteoporosis, increased bone fractures, skin thinning and skin bruising [7 -11] New developments in synthetic corticosteroids have lead to improvements in pulmonary targeting and reductions in adverse effects [12] Fate of ICS In order to produce further improvements in the area of ICS therapy it is necessary to be fully aware of the fate of corticosteroids following inhalation, as described in Figure 11. Following inhalation a fraction of drug is deposited in the lungs, a certain portion of the dose can be deposited in the oropharynx and if this is not rinsed may be sw allowed. Subsequently the swallowed portion is available for oral absorption; if the drug has significant oral bioavailability, it enters the systemic circulation to produce unwanted adverse effects. Drug may enter the systemic circulation as a result of absorption from the lung and gastrointestinal tract, both of which contribute towards systemic availability of the ICS [12 -14] Newer ICS such as fluticasone propionate (FP) and mometasone furoate (MF) have negligible oral bioavailabi lity (Foral), therefore this route does not significantly contribute to the systemically available dose [14] A fraction is deposited in the lungs where it must first undergo dissolution to exert its desired

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19 pharmacologic effect; diss olved drug will be absorbed into the systemic circulation, particulate matter deposited in the conducting airways however, will be subject to clearance by the mucociliary escalator. In order to achieve high pulmonary targeting drug dissolution should be s low, allowing pulmonary glucocorticoid receptors to be occupied for a longer period of time. On the other hand if particles dissolve much slower than the rate at which they are cleared from the lung, receptor occupancy over a given time period will be red uced. In order to achieve pulmonary targeting, corticosteroids should provide drug release at a rate comparable to their clearance from the lung [15] Pulmonary T argeting of ICS Receptor occupancy -time profiles may serve as a surrogate marker for pulmonary and systemic effects; cumulative receptor occupancy allows direct comparisons to be made between local efficacy and systemic adverse effects [16] For a given dose i ncreased cumulative receptor occupancy in the lungs compared to the systemic circulation results in improved pulmonary targeting; providing better efficacy and reduced systemic adverse effects [12, 14, 15] Various factor s are important for pulmonary targeting of ICS, these are based on pulmonary or systemic factors. Pulmonary pharmacokinetic factors include deposition efficiency and region of deposition in the lung of the ICS, also pulmonary residence time and lung tissue binding are important. Pulmonary pharmacodynamic factors include receptor binding affinity and selectivity. Systemic factors include oral bioavailability, clearance, volume of distribution, protein binding, tissue binding and affinity to transporters [15] To

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20 summarize, pulmonary targeting for inhaled drugs is influenced by both the inhaler device and formulation and are discussed below [17] Ideal C orticosteroid An ideal ICS should have high pulmonary targeting, defined as the difference between cumulative lung and liver receptor occupancies. High pulmonary versus oropharnygeal deposition is desired, however as a significant fraction of inhaled drug is deposited in the orophary n x from which it is subsequently swallowed, it follows that an ideal ICS should posses negligible oral bioavailability. This is the case for newer corticosteroids such as flutic asone propionate, mometasone furoate, ciclesonide and des -ciclesonide, which have less than 1% oral bioavailability [18, 19] Systemic bioavailability is the sum of pulmonary and oral bioavailability, si nce oral bioavailability is negligible for the newer corticosteroids, pulmonary absorption is responsible for the significant portion of drug available to the systemic circulation. An ideal corticosteroid should be removed rapidly from the systemic circul ation to reduce systemic adverse effects. The maximal clearance rate from the liver is approximately 90 L/min, which is equal to the liver blood flow, most corticosteroids posses s systemic clearance rates that are similar to blood flow. ICS have high hep atic extraction, resulting in low systemic exposure and enhanced pulmonary targeting, described in Table 1 1. Further improvements in this area will be difficult as clearance of ICS are already close to that of liver blood flow improvements may be achiev ed through the possibility of increased extra hepatic metabolism Increased pulmonary residence prolongs drug action in the lungs, thereby improving efficacy. Utilization of slow release inhaled formulations will increase pulmonary residence time and tar geting. In the case of slowly dissolving

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21 drugs the effect of mucociliary clearance predominates; as a result drug particles are removed from the lungs before they are able to completely dissolve, thus reducing pulmonary targeting [15] Inhaled formulations capable of providing slow release combined with reduced clearance from the lung are able to improve pulmonary targeting; this ap proach increases efficacy allowing lower doses to be administered. As a result fewer systemic side effects will be experienced. Inhaler D evice Commonly used inhaler devices include Dry Powder Inhalers (DPIs) and pressurized Metered Dose Inhalers (pMDIs) and to a lesser extent nebulizers in addition a new category of inhalers termed soft mist inhalers (SMI) has been developed recently pMDI devices utilize propellant s to aerosolize a solution or suspension, however they require the patient to co ordinate both device actuation and inspiration; resulting in poor inhaler technique especially in young children and the elderly. In addition to this a blast of high velocity cold propellant can impact the oropharynx in a small number of patients this will induce a gag reflex [20] DPIs avoid problems with inhaler technique as they are usually driven by the patients own breath. This itself however creates difficulty in many asthmatic and COPD patients who do not have the necessary lung function [12] The inhalation device plays a vital role in delivering drug to the lung. I n the early years of ICS therapy doses were delivered with low pulmonary deposition efficiencies of approximately 10% or less [12] In recent times however, new inhaler devices capable of delivering >30% of the d ose to the lung have been developed [13, 21, 22] An example of which comes from the re development of the Qvar containing beclomethasone dipropionate (BDP) from a CFC (CFC -BDP) to a HFA (HFABDP)

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22 propellant, as a result pulmonary deposition increased from <10 % to 5060 %. In addition HFA -BDP deposited less drug on the oropharynx and resulted in a more even distribution throughout the lung [21] Modulite technology was developed to allow H FA propellants to replace CFCs in pMDIs. The formoterol Modulite HFA inhaler provides a respirable fraction of 35% and 32.5% respectively. Slower plume velocity reduces impaction of the aerosol in the throat and therefore leads to greater lung depositio n. In addition slower plume velocities over a longer time period reduce problems associated with poor hand-breath coordination; the inhaled dose is delivered over a longer time period, therefore despite poor timing of the inhaler device with the patient s inspiration, majority o f the dose will still be inhaled High speed photography of the aerosol cloud produced following actuation of the Modulite device shows a greater plume length, reduced velocity and extended spray duration of up to 220 ms with for moterol Modulite compared to CFC propelled salmeterol. This approach leads to redu ced deposition in the oropharynx and subsequently improved pulmonary targeting [23] The Respimat belongs to a new class of inhaler devices termed soft mist inhaler s (SMIs) It does not contain any propellant instead it utilizes the mechanical energy from a spring to aerosolize droplets through a two channel nozzle, resulting in the production of an aerosol as a result of impaction of two converging jets of liquid at a carefully controlled angle. The uniblock is the key element of the Respimat, constructed with a silicon wafer bonded to a small ( 2 mm x 2.5 mm) b orosilicate glass plate. The spray is generated over approximately 1.5 seconds ; this results in a gentle mist of respirable particles and allows more time for the patient to coordinate device actuation with inspiration As the Respimat p roduces a gentle mist which emerges at

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23 a slower velocity compared to the aerosol cloud from pMDIS, less drug is deposited in the oropharynx [22, 24, 25] T he Respimat delivers a greater fraction to the lung, it is p ossible to reduce the Metered Dose (MD) and still maintain the same c linical efficacy. The dose of i pratropium/fenoterol hydrobromide delivered using the Respimat may be reduced by 50% of the dose administered using a pMDI with a spacer device, while producing the same clinical effect [25] Improvements in pulmonary deposition may increase pulmonary selectivity, especially for corticosteroids with high oral bioavailability Improvements in lung deposition are not as significant for newer ICS with low oral bioavailability [12] Influence of Mass Median Aerodynamic Diameter Deposition of particles in the lung is governed by its mass median aerodynamic diameter (MMAD). Respirable particles are in the range 1 5m and will deposit effectively in the lung. Very small particles (<1m) will not be effectively deposited and are exhaled. Large particles (>10m) will be deposited in the tracheobronchial region and then swallowed [26] Depending on the oral bioavailability of the compound this may contribute to therapeutic efficacy and adverse effects. For example many of the more recently developed inhaled corticosteroids such as fluticasone propionate, mometasone furoate, ciclesonide and des -ciclesonide have minimal oral bioavailability thus this route does not contribute significantly towards systemic drug exposure [12] Controlled Release Inhaled F ormulations A number of methods to prolong pulmonary residence time of inhaled formulations have been employed. For example budesonide forms reversible fatty esters within cells, this forms a depot of inactive drug until the ester is broken down to the active form. This fatty acid esterification of budesonide, prolongs its action within the lungs

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24 allowing once daily dosing and reduced systemic adverse effects [27 31] In order for a corticosteroid to undergo fatty acid esterification it must possess a steric -hindrance -free hydroxyl group at the carbon 21 position [31] Liposomal formulations have also been investigated to provide slow release follow ing inhalation, however, liposo mes have problems with stability due to leakage of drug duri ng storage or jet -milling of lyophilized formulations [16, 32] Solid lipid nanoparticles and microparticles provide slow drug release and increased stability in comparison to liposomes [33 35] Biodegradable polymers such as PLGA and PLA are used widely in the drug delivery due to their biocompatible nature and approval for use as excipients [36] PLGA microspheres have been formulated to release dexamethasone continuously over one month [37] Drug particles may also be coated with nano thin layers of polymer to produce slow release formulations, pulsed laser deposition of PLA onto glucocorticoids resulted in slower release budesonide [38] Large porous low density microparticles containing albuterol demonstrated sustained bronchodilation over at least 16 hours compared to 5 hours provided by nonporous particles of similar MMAD [39] PLGA and PLA nanoparticles have received a great deal of attention, a number of studies have described their use for providing slow release. Usually a biphasic release profile was observed in most cases, initially a burst release is observed, followed by a slow release profile [40 -46] Slow release formulation of nanoparticles may be problematic, due to their large surface area over which diffusion out of the polymer matrix occurs many studies observed very fast drug release [47 -49] Mucoadhesive properties of chitosan allow nanoparticle retenti on time to be further increased, a study by Yamamoto et al showed surface modified PLGA

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25 nanoparticles, modified with chitosan had prolonged effects compared with unmodified particles and also had reduced clearance[50] Polymeric N anoparticles Polymer nanoparticles of PLGA or PLA are increasingly becoming the focus of attention as they are biocompatible and biodegradable [36, 5 1] Degradation products are glycolic and lactic acid, natural bi -products of the Krebs cycle, readily eliminated from the body by further breakdown to carbon dioxide and water [36, 51] PLGA degradation occurs by hydrolysis, and is dependent on molecular weight, conformation and polymer composition [52, 53] Rates of polymer degradation are fastest when the composition is 50% lactic acid and 50% glycolic acid, thus we will use this for nanoparticle production [53] PLGA degradation however, may also be dependent on the type of drug encapsulated [54] Polymeric corticosteroid nanoparticles have been developed for a number of medical applications, for example, cancer, arthritis, choroidal neovascula rization and neural drug delivery to name a few [40, 42, 43, 55-59] Techniques for production of polymeric nanoparticles include solvent evaporation, nanoprecipitation, supercritical fluid prec ipitation, wet milling and high pressure homogenization [60 64] The solvent evaporation technique is used commonly, it involves the production of a microemulsion in which both drug and polymer are dissolved in the organic phase, F igure 21 shows the scheme of production of nanoparticles using the solvent evaporation technique. The organic phase is usually a volatile compound such as dichloromethane or acetone, this will diffuse into the aqueous phase and evaporate, leaving behind a nanoparticle suspension. Budhian et al explored the influence of various production parameters on particle size and drug encapsulation efficiency [65]

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26 Increased polymer lead to improved drug encapsulation, with a gradual increase in particle size. PVA surfactant concentration less than 0.5% w/v resulted in a bimodal particle size distributi on due to insufficient stability of the microemulsion. Solid Lipid Nanoparticles (SLN) and M icroparticles Solid lipid nanoparticles (SLN) have been studied for their potential in providing sustained drug delivery [34] Dexamethasone SLN particles have been developed using soybean lecithin and gl ycerol tristearate. A biphasic in vitro release was observed for these particles with an initial burst of approximately 70% followed by slower release [35] The lipid matrix of these SLN dispersion is more mobile in comparison to PLGA nanoparticles, thus slow release drug formulations may be more challenging [34] Slow release salbutamol acetonide lipid micro particles have also been developed by Jaspart et al. Encapsulation efficiency of the f ormulations developed was greater than 87%, however high drug loading of around 25% w/w resulted in crystallization of drug particles outside of the micro particles In vitro release studies determined micro particles formulated with lower drug loading prod uced a slower drug release, however all solid lipid micro particles released at a slower rate compared to pure salbutamol acetonide [33] Low Density Microspheres Large porous particles are more efficien tly aerosolized as they produce fewer aggregates and are easier to re disperse within an air stream. It is also possible that these large porous microparticles are too large to be engulfed by macrophages [66] Large porous microparticles can be produced by spray drying, this results in particles with a similar MMAD to smaller non porous particles. Large porous e stradiol particles were aerosolized into the lungs of rats using an endotracheal tube. These particles

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27 were shown to release over a longer time period of approximately 5 days compared to only release over 1 day with nonporous particles of similar MMAD [67] Pulsed Laser D eposition (PLD) A nano -thin coating of a polymer such as PLA or PLGA may be applied with the use of Pulsed Laser Deposition (PLD). High energy pulses of ultraviolet light are directed onto a polymer disc to create a plume of nanoparticles which are subsequently deposited onto a dry powder, thereby producing a nano -thin coating as shown in Figure 1 -2 [68] PLD can be used to produce sustained drug release, mean dissolution time for budesonide particles coated was 4.7 0.1 hours compared to only 1.2 0.5 hours for uncoated budesonide. Improved pulmonary targeting of PLA coated budesonide was demonstrated in neonatal rats. Table 1 2 shows AUC of receptor occupancy -time profiles of budesonide in neonatal rats in the lung and liver. It was seen that following intratracheal administration of uncoated budesonide, AUC for the lung and liver were indistinguishable from one another, whereas following administration of PLA coated budesonide a higher AUClung was seen compared to AUCliver; indicating improvement in pulmonary targeting with PLA coated budesonide [69] Physicochemical diff erences between drugs also play an important role in influencing improvements in pulmonary targeting following polymer coating. Triamcinolone Acetonide (TA) and budesonide were coated with PLA, no difference between dissolution profiles was observed between coated and uncoated TA, however the more lipophi lic corticosteroid, budesonide clearly show ed slower drug release after coating with PLA [70]

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28 Oligolactic Acid (OLA) Oligolactic acids (OLAs) are short chain versions of polylactides, which are biocompatible and approved by the FDA for use with implantable devices. OLAs are formulated as excipients in MDI inhalers b y 3M, they have been used as solubilizers, suspending agents and produce sustained release within the lung. Sustained release is achieved by the formation of OLA -drug matrix (solution formulations) or OLA -coated drug particle (suspension formulations) [71, 72] Pro -drugs Many ICS, such as BUD or FP are administered in their pharmacologically active form; others may be inhaled as prodrugs that must first undergo conversion to their active form. This approach can be utilized to improve pu lmonary by achieving therapeutic drug concentrations at the target site whilst minimizing unwanted side effects at other sites. Deposition of ICS in the mouth and oropharynx can give rise to adverse effects such as oral candidiasis and dysphonia. Two currently available ICS include beclomethasone dipropionate and ciclesonide are converted to their active metabolites by esterase enzymes within the pulmonary epithelium [73 -75] It has been shown that bioactivation of ciclesonide is very low in the oropharynx, therefore less acti ve drug is present in comparison administration of budesonide or FP [14] Ciclesonide, a newer inhaled corticosteroid is converted to its active metabolite by ester cleavage at the C21 position, resulting in the formation of desisobutyryl -ciclesonide with 100-fold higher potency. Combined with high systemic clearance of both the parent compound and the active metabolite, local and systemic adverse effects are reduced. In addition the active metabolite of ciclesonide undergoes fatty acid esterification within the lung, to prolong retention within the lung, described in further detail later [74 77]

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29 Ester F ormation Pulmonary retention of inhaled corticosteroids may be enhanced due to reversibl e fatty acid esterification. Conjugation has only been reported for budesonide, triamcinolone acetonide and des -ciclesonide. This process occurs within cells in the lung and forms a reservoir of lipid conjugated drug. Conjugated corticosteroid is slowly hydrolyzed by enzymes to release free drug. In order for this fatty acid conjugation to take place the ICS must possess a steric hindrance-f re e hydroxyl group at the carbon 21 position. Reversible fatty acid conjugation provides increased anti -inflammat ory action at the target site resulting in improved pulmonary selectivity [27 -31, 74, 75, 78 80] In general, rapid ester formation and slow ester hydrolysis leads to improved pulmonary targeting; this is further improved in combination with a high systemic clearance. The corticosteroid budesonide is moderately lipophilic with a relatively fast dis solution, followed by a rapid absorption from the lung. Budesonide would not be expected to have a prolonged duration of action in the lung when compared to fluticasone propionate, a corticosteroid with both higher lipophilicity and relative receptor affi nity [80] Format ion of the fatty acid ester, budesonide oleate however, prolongs retention in the airways and allows for once daily dosing [29] Ciclesonide is a corticosteroid which first undergoes metabolism to the act ive form, des -ciclesonide followed by reversible fatty acid ester formation. This ester formation was confirmed to occur in the human lung in a single dose, open -label, non-randomized study in 20 patients. The metabolites des -ciclesonide, des -ciclesonide oleate and des ciclesonide palmitate were detected in the central and peripheral lung tissue [73]

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30 Budesonide, however was shown to be esterified more rapidly and to a greater extent in comparison to ciclesonide in rat tracheal tissue [79] Future developments in corticosteroids may involve selection of compounds which undergo fatty acid es terification [81] Liposomes Liposomes have been extensively studied due to their ability to incorporate both hydrophilic and hydrophobic drugs, as well as ability to produce a varie ty of particle sizes and have been used as drug carriers since the late 1960s [82] Liposomes are composed of phospholipids which are endogenous in the lung, improving compatibility. Nebulized BDP -DLPC liposomes administered to healthy volunteers are w ell tolerated i n doses equivalent to those currently used for treatment in asthma [83] Modified liposomes may be used to target delivery to different cells, liposomes prepared with mannosylated cholesterol derivatives can enhance uptake into macrophage cells [84] Liposomal formulations have also been investigated to provide slow release following inhalation, however, liposomes have problems with stability due to leakage of drug during s torage or jet -milling of lyophilized formulations [16, 32] Solid lipid nanoparticles and microparticles provide slow drug release and increased stability in comparison to liposomes [33 -35] It has been demonstrated that slow drug release from liposomes is able to improve pulmonary selectivity in the rat model. Triamcinolone Acetonide Phosphate (TAP) liposomes sized 200nm and 800nm were shown to produce biphasic drug release in vitro with the later resulting in slower release ). In vivo studies in male F -344 rats were performed in order to compare pulmonary selectivity of these liposomes with TA solution and liposomes. TA liposomes release drug rapidly under sink conditions, and thus both the TA solution and liposome formulation would not be

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31 expected t o produce pulmonary targeting, Table 1-3 show 800nm TAP liposomes produced the greatest pulmonary targeting; pulmonary targeting is defined as the area under the curve of the receptor occupancy time profile of the lung compared to the liver (Pulmonary targeting (%*hr) = AUClung AUCliver). Slow drug release from liposomes allows reduction in dose and dosing frequency, thereby reducing systemic adverse effects observed whilst maintaining clin ical effects in the lung [16] It has been shown that budesonide encapsulated into stealth liposomes, delivered once a week was able to provide an equivalent anti -inflammatory e ffect as once daily administration of budesonide [85] Enhancing Mucoadhesive P roperties Chitosan is a polysaccharide containing an amino group which may be positively charg ed, particles composed or coated with chitosan increase interaction with negatively charged lung epithelial cells. In addition chitosan may also influence release characteristics of drug particles. Gelatin has also been utilized to produce particles. Ri fampicin liposomes coated with chitosan had greater mucoadhesive properties and lower toxicity towards A549 epithelial cells compared to uncoated liposomes. Table 1 4 demonstrates the relationship between zeta potential and mucoadhesive properties of rifa mpicin liposomes. Uncoated negatively charged liposomes showed the lowest amount of adhesion, followed by uncharged liposomes, with positively charged chitosan coated liposomes showing the most mucoadhesion [86] Objectives Increased pulmonary retention has been observed following inhalation of nanometer sized particles in comparison to micronized particles [87, 88] In addition polymers such as PLGA and PLA are able to reduce drug release from nanoparticles.

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32 We hypothesize that nanoparticles may be spray dried with lactose to produce nanocomposite microspheres; these microspheres will be ca pable of delivery to the lung following inhalation. The above hypotheses will be tested by the following specific aims; Preparation and characterization of slow release polymeric corticosteroid nanoparticles. In vitro drug release testing to determine sl ow release characteristics of nanoparticles. Design of a spray dryer able to enable nanoparticles to be spray dried at temperatures lower than operating temperatures of commercially available spray dryers. Preparation of spray dried nanoparticles to form n anocomposite microspheres with MMAD in the respirable range. Determine in vitro release characteristics of spray dried formulations to investigate the effect of spray drying on release rate

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33 Figure 1 1 Fate of Inhaled Corticosteroids [12]

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34 Table 1 1 PKPD properties of inhaled corticosteroids [15] ICS RRA CL (L/hr) Vdss (L) F oral (%) f u (%) BDP 53 150 20 15 20 13 BMP 1022 120* 424 26 FLU 190 57 96 20 20 TA 233 37 103 23 29 BUD 935 84 18,311 12 FP 1800 69 318 <1 10 MF 2900 54 <1 1 2 CIC 12 140 207 <1 1 Des CIC 1200 228* 897 <1 <1

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35 Figure 12 Pulsed laser deposition (PLD) Nanoclusters of polymer from the target are deposited on larger micronized drug particles as continuous coatings that sustain the release rate of drug in solution [38]

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36 Table 1 2 Average AUCs (n=3) in the lung, liver and brain and pulmonary targeting (PT) in neonatal rats after intratracheal administration (50 g/kg) of uncoated budesonide and PLA coated budesonide [69] Formulation Dose (g/kg) AUC l ung AUC l iver AUC b rain PT (AUClung/AUCliver) Uncoated budesonide 50 58.4 12.9 56.4 6.8 38.3 6.7 1.03 0.13 PLA coated budesonide 50 75.8 3.7 46.6 14.5 29 7 1.72 0.46

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37 Table 1 3 Cumulative Receptor Occupancy (AUC), Pulmonary Targeting and Mean Pulmonary Effect Times (MET) After Intratracheal Administration of Escalating Doses of TAP in 800 nm Liposomes [16] TAP sol TA lip 200 nm TAP lip 200 nm TAP lip 800 nm Lung 370 50 320 85 770 120 1070 70 Liver 340 40 380 10 620 150 700 140 Pulmonary Targeting (%*hr)(AUClungAUCliver) 30 10 -60 80 150 60 370 70 Pulmonary Targeting (%*hr)(AUClung/AUCliver) 1.0 0.85 1.2 1.5 Mean Pulmonary Effect Time (hr) 3 2.4 5.7 >6.2

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38 Table 1 4 Influence of liposome composition on mucoadhesion and zeta potential [86] L iposome composition Mucin adsorbed on liposomes (%) Mean value (SD) Zeta Potential (mV) PC/Chol 17.0 8.3 +0.09 0.54 PC/PG/Chol 7.4 4.4 22.9 2.1 [PC/Chol] CHT 47.1 1.2 +4.4 1.9 [PC/PG/Chol] CHT 90.9 7.6 +24.98 0.91 DSPC/Chol 46.2 4.1 +0.93 0.77 DSPC/PG/Chol (9:1:5) 25.1 8.1 19.9 2.3 [DSPC/Chol] CHT 66.6 2.2 +5.4 2.7 [DSCP/PG/Chol] CHT 93.1 4.1 24.43 0.62

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39 CHAPTER 2 DEVELOPMENT AND CHARACTERIZATION OF SLOW RELEASE POLYMERIC CORTICOSTEROID NANOP ARTICLES Introduction Nanoparticles may have the possibility of providing increased retention in the airways compared to micron sized particles [87] Nanoparticles have been extensively studied in the field of toxicology. A number o f studies have been conducted in order to determine deposition, retention and translocation of ultrafine particles in the lung. Technetium Tc 99m (99mTc) -r adiolabeled 100nm carbon particles were administered to healthy subjects and COPD patients by nebulizer. The central / peripheral (C/P) distribution was controlled by administering either a shallow or deep aerosol bolus A shallow aerosol bolus is used to deliver either more centrally compared to a deep aerosol bolus inhalation. 48 hours following a shallow aerosol bolus, 70% and 82% of particles were retained within the airways of healthy non-smokers and COPD patients respectively It has been shown that nanoparticles are retained in the lung for a longer time period compared with micron sized particles; 24 hours following inhalation >70% of nanoparticles are retained in the airways, in comparison only 10% of particles greater than 6 m are retained [87, 88] It has been proposed that increas ed retention on nanoparticles in the lung is as a result of greater displacement of nanoparticles into the aqueous surfactant film compared to micron sized particles [87, 89, 90] In recent times nanotechnology has received much attention, potential applications include imaging and diagnosis, targeted drug delivery and controlled drug release [91, 92] Nanotechnology itself is not a new concept, liposomes were first used as drug carriers in the late 1960s [82] Polymer nanoparticles of PLGA or PLA are increasingly becoming the focus of attenti on as they are biocompatible, biodegradable and have

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40 been approved by the FDA for use in implantable devices [36, 51] Degradation products are gly colic and lactic acid, natural bi products of the Krebs cycle, readily eliminated from the body by further breakdown to carbon dioxide and water [36, 51] PLGA degradation occurs by hydrolysis, and is dependent on molecular weight, conformation and polymer composition [52, 53] Rates of polymer degradation are fastest when the composition is 50% lactic acid and 50% glycolic acid, thus this was used for production of nanoparticle s [53] PLGA degradation however, may also be dependent on the type of drug encapsulated [54] Drug release from particles is also influenced by polymer composition. Increased lactide:glycolide results in slow er release, possibly due to increased hydrophobicity of the polymer as well as increased solid state solubility of hydrophobic drugs in the polymer matrix It has also been determined that PLGA with an ester terminated end group released hydrophobic drugs at a slower release rate compared with a carboxylic acid end group [93] Polymeric corticosteroid nanoparticles have been developed for a number of medical applications, for example, cancer, arthritis, choroidal neovascularization and neural drug delivery to name a few [40, 42, 43, 55-59] Techniques for production of polymeric nanoparticles include solvent evaporation, nanoprecipitation, supercritical fluid precipitation, wet milling and high pressure homogenization [60 64] Advantages and disadvantages of the various methods for nanoparticle production are discussed in Table 2 1. The solvent evaporation technique is used in the production of monodispersed spherical polymeric nanoparticles. This technique is used commonly, it involves the production of a microemulsion in which both drug and polymer are dissolved in the

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41 organic phase, F igure 21 shows the scheme of production of nanoparticles using the solvent evaporation technique. The organic phase is usually a volatile compound such as dichloromethane or acetone, this will diffuse into the aqueous phase and evaporate, leaving behind a nanoparticle suspension. Nanoparticles may be collected and washed by centrifugation, followed by lyophilization to allow storage as a dry powder. A number of process parameters may be manipulated to influence particle size and drug loading. Budhian et al explored the influence of various production parameters on particle size and drug encapsulation efficiency [65] Increased polymer content lead to improved drug encapsulation, with a gradual increase in particle size. Increased polymer content allows more drug to be dispersed within the polymer matrix, thus resulting in higher encapsulation e fficiency. As the polymer contained in the organic phase increases however, viscosity of this phase is also increased, thus resulting in formation of larger o/w microemulsion droplets and larger nano particle size. PVA surfactant concentration less than 0 .5% w/v resulted in a bimodal particle size distribution due to insufficient surfactant concentration for production of a stable microemulsion. The solvent evaporation method has been used to produce slow release nanoparticles of a number of different comp ounds including paclitaxel, etanidazole, bezopsoralen, flurbiprofen and dexamethasone; a biphasic in vitro release was observed in most cases [40 -46] It is difficult to compare results from different studies however, as drug release rates depend on release conditions such as presence of sink conditions, release media and sample separation methods (filtration, centrifugation or dialysis). Though each method has its advantages it is possible that certain artifacts may result due to the separation method used. For example retardation of drug release

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42 from dialysis bags may cause release profiles to reflect diffus ion through the dialysis membrane rather than from the formulation [34, 94] In these situations, it is possible for the drug to interact with the dialysis membrane, slow release observed is due to an interaction with the membrane rather than the actual release rate from the nanoparticl e formulation; therefore it is necessary to ensure that the chosen dialysis membrane does not interact with the drug of interest, as well as inclusion of adequate controls in the experiment. In addition, release rate can further be reduced by coating with another polymer such as chitosan. Chitosan is polysaccharide derived from the deacetylation of chitin, commonly obtained from the shells of crustaceans. It is biocompatible and can be degraded by lysozymes present in all mammalian cells [95] Cationic chitosan is able to coat nanoparticles due to an electrostatic interaction with the negatively charged carboxylic acid end group present on the polymer PLGA or PLA [96] Chitosan has been used to coat PLGA nanoparticl es as well as liposomes [86, 9699] PL GA nanoparticles surface modified with chitosan also had increased retention in the lung compared to PLGA nanoparticles due to improved mucoadhesive properties of chitosan [50] Hypothesis We hypothesize that polymeric nanoparticles with a monodispersed particle size distribution can be formed using the solvent evaporation method. Various parameters during this process will be manipulated to influence particle size and drug encapsulation. We hypothesize that increasing polymer content will increase particle size and drug encapsulation efficiency. Also increasing PVA concentration will allow smaller oil in water microemulsion droplets to be formed, resulting in smaller particles with lower drug encapsulation. Larger particles may be formed with reducing PVA

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43 concentration, however a minimum content must be achieved in order to sufficiently stabilize the microemulsion and result in a monodispersed particle size distribution. It is hypothesize d that c hitosan coated nanoparticles will be able to release at a slower rate in comparison with uncoated particles. Electrostatic interaction between the positively charged chitosan and negatively charged polymer allows chitosan to coat polymeric nanopar ticles It is important that the polymer chain is terminated with a carboxylic acid end group and is not ester terminated, to ensure interaction with the chitosan. Methods to produce chitosan coated particles will be investigated and optimized. In vitro release testing will be performed by the batch/filter method [100] This method reduces artifacts which may be observed by using the dialysis method for separation of particles from free drug. It is hypothesized that polymeric nanoparticles will release drug at a slower rate compared to the control micronized drug. Materials and Methods Chemicals Micronized TA was purchased from PCCA Inc. (Houston, TX, USA). MF was donated by Ipca laboratorie s ltd (Mumbai, India). PLGA and PLA were purchased from Lactel Absorbable Polymers (Pelham, AL, USA). DCM and ACN were purchased from Fisher Scientific (Pittsburgh, PA, USA). PVA and medium molecular weight chitosan w ere obtained from Sigma Chemical Co. (St. Louis, MO). Nanoparticle Preparation Polymeric nanoparticles of either TA or MF were prepared using the solvent evaporation technique [47] TA will be used as a model drug to allow investigation of

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44 process parameters on particle size, drug encapsulation efficiency and in vitro release profiles. The polymer used was either DL -PLGA (50:50 inherent (0.55 -0.75 dL/g inherent viscosity) or DL -PLA (0.54 dL/g inherent viscosity). In brief 10 mg of drug and 200-400 mg polymer were dissolved in 5 mL DCM, this was preemulsified with 5 mL of the aqueous PVA phase by vortexing for 30 seconds in a 20 mL glass scintillation vial using the Fisher vortex genie 2 at speed 9. The pre emulsion was added to the remaining 45 mL PVA solution and sonicated on ice (Sonics Vibra Cell Ultrasonicator Newtown, CT) at 60 W for 5 minutes to form an oil in water microemulsion. PVA concentration was varied between 1 and 3% w/v for each PLGA level, (n=3) DCM was allowed to evaporate under gentle stirring on the magnetic stirrer for 4 hours to produce a nanoparticle suspension which was collected by centrifugation at 20,000 rpm for 40 minutes and washed Beckman J221 (Beckman Coulter, Inc.Fullerton, CA) using the JA -20 rotor Nanoparticles were lypophilized using a Labconco freeze dryer (Labconco Corporation, Kansas City, MO) Dried nanoparticles were stored in amber glass vials at 4 C in a dessi cator. Two different methods to prepare chitosan coated PLA nanoparticles were studied. Figure 2 -2 shows details of these two methods, the incubation method involves dispersion of PLA nanoparticles in either 0.1% or 1% w/v chitosan solution. In situ coat ing allows the chitosan coating to be applied in the same step as formation of the nanoparticles. Nanoparticles are prepared using the solvent evaporation technique, however chitosan is already dissolved in the aqueous PVA phase of the microemulsion. Dur ing formation of the nanoparticle suspension, the positively charged chitosan is coated onto the negative PLA particles by an electrostatic interaction. The optimal

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45 method was chosen based on the ability to for m unimodal nanoparticles. The aqueous phase of the oil in water microemulsion was produced using PVA and chitosan in 1% v/v acetic acid in DDW. 400 mg PLA and 10 mg MF were dissolved into 5 mL DCM, this was preemulsified with 5 mL of the aqueous phase by vortexing for 30 seconds. This was combined with the remaining 45 mL of the aqueous phase and sonicated on ice for 5 minutes at 60 W to produce an oil in water microemulsion. The emulsion was gently stirred for a further 4 hours to allow the DCM to evaporate and leave behind a chitosan coated nan oparticle suspension. The resultant suspension was centrifuged and washed using a 1% v/v acetic acid solution. Chitosan coated nanoparticles were dispersed in a small volume of double distilled water followed by lyophiliz ation and stored in amber glass v ials in a desiccator at 4 C Drug loading and Encapsulation E fficiency In order to determine the drug content of the lyophilized nanoparticles approximately 5 mg of the dry powder was weighed. The particles were then dissolved using 2 mL DCM and placed on an orbital shaker (Bellco biotechnology, Vineland, NJ) overnight to en sure disintegration of the nanoparticles. The DCM was evaporated off using a Jouan RC10.10 vacuum centrifuge (Thermal Fisher, Asheville, NC), the dried residue was dissolved in 1 mL mobile phase (ACN:DDW, 60:40) and analyzed by HPLC -UV (Hewlett Packard Se ries 1050) using a Phenomenex Ultracarb 30 4.6x150 mm ODS column using mobile phase at a flow rate of 1.2 mL/min sample peaks were quantified at 254 nm wavelength. Concentrations for the caliberation curve were 10, 20, 40, 60, 80 and 100 g/mL with an R s quared value of at least 0.997. Figure 23 describes the method by which EE and drug loading of nanoparticles was determined.

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46 Drug loading and encapsulation efficiency were determined by the equations below. Theoretical drug loading (% w/w) was calculat ed by calculating how much drug would be theoretically present in 100 g of the lyophilized powder provided that no drug or polymer is lost during the nanoparticle production process, the equation for which is described below. ( % / ) = 100 ( % / ) = + 100 ( % / ) = ( % / ) ( % / ) 100 Particle S ize Z eta P otential and M orphology Particle size and zeta potential were determined using the Nanotrac (Microtrac) and Brookhaven ZetaPlus. SEM was performed to observe particle morphology using scanning electron microscope (SEM) JEOL JSM 6335F instrument (Major Analytical Instrument Center (MAIC), UF, Gainesville, FL). Briefly, formulations were placed on carbon stubs which were coated with c arbon using a vacuum evaporator SEM was conducted using 2 kV. In vitro D rug R elease Study In vitro release was performed in 10 0 mL 1% v/v tween 80 in PBS, shaken at 30 rpm in a hot shaker (Bellco biotechnology, Vineland, NJ) at 37 C over 24 hours under constant sink conditions. Tween 80 was used to increase the saturation concentration of TA and MF in the dissolution media to 83 g/ml and 25 g/ml with addition of 1% v/v tween 80 at 37 C. In order to maintain sink conditions maximal concentrations obtained did not exceed 10% of the saturation concentration. At specific time points (0, 15, 30,

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47 90, 120, 180, 240, 360, 600 and 1440 minutes), 1 mL was removed using a pipette and subsequently filtered using a 0.02 m filter (Whatman Anotop plus filter), 1 mL fresh buffer was replaced into the dissolution media. The maximum concentration obtained was determined based on analysis of remaining undissolved drug at the end of the dissolution testing. A fter the last time point, the dissolution media was centrifuged at 20,000 rpm for 40 mins using the Beckman J221 (Beckman Coulter, Inc.Fullerton, CA) using the JA -20 rotor. The pellet was dissolved in 2 mL DCM and placed on an orbital shaker overnight (Bellc o biotechnology, Vineland, NJ) to ensure disintegration of the particles. HPLC was used to quantify released drug concentrations using a caliberation curve ranging from 0.5 -10 ug/mL. Controls used were micronized TA and MF contained in the Asmanex formu lation from the reservoir based inhaler device, Twi st halerTM, Schering -Plough. Results and Discussion Influence of PLGA and PVA on Encapsulation Efficiency and Particle Size of Nanoparticles A factorial design was implemented to determine influence of PLGA and PVA concentration on particle size and encapsulation of TA within nanoparticles. Encapsulation efficiency of TA nanoparticles decreased with increas ing PVA concentration, shown in Figure 2-4 as expected. Increased PVA surfactant concentration in the microemulsion aqueous phase results in an increased solubility of TA in the aqueous phase. A consequence of this is increased loss of drug during nanoparticle production is reduced encapsulation efficiency and drug loading. Increasing PLGA content did not have any significant effect on encapsulation efficiency. Increasing PVA concentrations resulted in a trend of decreased particle size, as shown

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48 in Figure 2 -5 Increased surfactant concentration in the oil -in water microemu lsion leads to the formation of smaller oil phase droplets; as a result particle size of nanoparticles decreases with increased PVA concentration. Increasing PLGA content results in the formation of larger nanoparticles, this is as a result of increased v iscosity of the oil phase of the microemulsion; larger organic phase droplets are formed leading to an increase in particle size. This is in agreement with studies performed by Budhian et al; increasing PVA concentration results in reduced drug loading in itially, between 12% w/v PVA then plateau. Particle size also decreased with increasing PVA concentration in their study, however with the use of >5% w/v PVA particle size increased. This is due to competing effects of increased stabilization of the mic roemulsion with increasing surfactant, which reduces particle size. High PVA concentration increases viscosity of the aqueous phase, leading to reduced net shear stress for droplet breakdown during formation of the microemulsion, therefore resulting in la rger particles [65] The PVA concentration range we investigated was 13% w/v, thus it was determined i ncreasing PVA concentration reduced encapsulation efficiency and a trend of reduced particle size was observed. Budhian et al also investigated the influence of PLGA content on particle size and drug loading. Increasing PLGA content lead to a gradual inc rease in particle size and drug loading, as was observed in our experiments producing TA -PLGA nanoparticles Influence of PLGA and PVA on Morphology of Nanoparticles TA crystals were observed for two formulations of TA -PLGA nanoparticles, these were both formulated using 200mg PLGA us ing 1% and 2% w/v PVA, SEM of these formulations are shown in F igure 2. 6 and F igure 2. 7 These TA crystals were only

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49 observed when particles were formulated with both low polymer contents and surfactant concentrations. There must be sufficient polymer for the drug to disperse within in the solid state; drug not present in the polymer matrix must then be washed off [93] When only 200 mg PLGA is used, there is insufficient polymer for the TA to disperse within, coupled with low PVA surfactant, unencapsulated drug is not washed away leaving behind free TA crystals. Although it appears that encapsulation efficiency for these two samples is greater than 50%, this is not due to TA trapped in PLGA nanoparticle matrix Figure 2. 8 shows spherical TA -PLGA nanoparticles produced with 400 mg PLGA and 1 % w/v PVA, no TA crystals were observed in this formulation. Formulations produced using 300 -400 mg PLGA did not contain any unencapsulated drug crystals, as determined by SEM. In Vitro Release Study of TA PLGA Nanoparticles Various methods to determine in vitro drug release of nanoparticles have been reported in the literature. Regenerated cellulose ester membranes and dialysis bags have been used to physically separate nanoparticles from the dissolution media. Preliminary studies performed however, indicated interaction of corticosteroids with the membrane and delayed release of TA from within dialysis bags. If this method is used, it is essential to find dialysis membranes which do not interact with the drug being studied, as well as use of adequate controls [94] Separation by filtration was used to determine in vitro release to avoid problems which occurred with the use of regenerated cellulose dialysis membranes. In vitro release studies were performed on TA -PLGA nanoparticles formulated with either 300 mg or 400 mg PLGA using both 1% and 2% w/v PVA. These formulations had good drug encapsulation ef ficiency as well as a unimodal particle size.

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50 In comparison to the control of mic ronized TA, nanoparticle formulations made with only 300mg PLGA released at a slightly faster rate, as shown in F igure 29 TA -PLGA nanoparticles formulated with 400mg PLGA were shown to release at the same rate as micronized TA, refer to F igure 2-1 0 Al l formulations tested release drug at a very similar rate to the control micronized TA. The surface area to mass ratio of nanoparticles is very high compared with micron sized particles, thus if no modification is made to the formulation, nanoparticles would release drug at a much faster rate compared to their micron sized counterpart. TA -PLGA nanoparticles posses some slow release characteristics as they are able to release drug at a similar rate to micronized TA. Chitosan Coated and Uncoated PLA MF Nano particles Two methods for production of chitosan coated polymeric nanoparticles were investigated. The chitosan coating procedure was optimized to produce particles coated particles with a unimodal particle size distribution. Lyophilized MF -PLA nanoparti cles that were incubated with either 0.1% or 1% w/v chitosan solutions resulted in nanoparticles that had multimodal particle size distributions. A possible explanation for this is due to incomplete dispersion of nanoparticles in the viscous chitosan solution, leading to coating of agglomerat ed nanoparticles SEM of MF -PLA nanoparticles incubated with a chitosan solution showed a bimodal particle size distribution (Figure 2 -1 1 ). In situ chitosan coating involves coating during the formation of nanoparticles. Chitosan is present in the aqueous phase of the microemulsion, positively charged chitosan adheres onto the negatively charged PLA nanoparticles. Addition of 1% w/v chitosan to the aqueous phase results in bimodal particle size distribution. High viscosity of 1% w/v chitosan solution resists droplet breakdown during

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51 sonication, producing some larger particles. In addition, particles produced from 1% w/v chitosan solutions are more difficult to collect by centrifugation, due to the high viscosity of the solution. The use of 0.1% w/v chitosan with in the situ coating procedure resulted in a unimodal particle distribution and can be seen by SEM (Figure 21 3 ) and can be comp ared to uncoated MF -PLA nanoparticles (Figure 21 2 ). Particle size of chitosan coated particles was 226 nm compared to 212 nm for uncoated particles. Although particle size does not significantly increase following coating with chitosan, the presence of chitosan is confirmed by the reversal of zeta potential from negative charge on uncoated nanoparticles to positive charge on coated particles, as shown in T able 2 3 In order for chitosan to coat particles, the PLA polymer must terminate with a carboxylic acid group and the amine group on chitosan must be protonated, with 1% v/v acetic acid. In vitro release of chitosan coated particles was compared to uncoated MF -PLA nanoparticles and Asmanex as a control are shown in F igure 21 4 A bipha sic release was observed for both the chitosan coated and uncoated nanoparticles. Ini tially the burst release of MF -PLA nanoparticles is similar to the Asmanex formulation followed by slo w release. Chitosan coated MF -PLA nanoparticles (CH MF) release a pproximately 20% by 30 minutes, followed by a much slower release rate compared to MF or Asmanex After 24 hours of in vitro release, CH -MF nanoparticles only released 43% MF; in comparison, all MF from the Asmanex formulation was released in 1 hour. In vitro release data show that chitosan coating allows much slower drug release from nanoparticles. Drug r elease occurs by diffusion out of the nanoparticles as well as due to degradation of chitosan and PLA In the short term diffusion processes

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52 deter mine drug release, however over longer time periods degradation of chitosan and PLA play a more dominant role. Similar observations were made from testosterone contained in a PLGA film, initial release was as a result of diffusion from the film, followed by a slower release due to the hydrolytic degradation of PLGA [101] Both chitosan coated and uncoated nanoparticles formulated with MF were able to release at a slower rate compared to the formulation contained within the commercially available Asmanex TwistnalerTM, currently used by oral inhalation for longterm asthma management. PLGA microspheres of rifampicin coated with chitosan were shown to produce a smaller burst effect compared to uncoated particles by Manca et al [97] these results are in agreement with our observations from chitosan coated and uncoated PLA nanoparticles. Conclusion Increased PLGA content resulted in a general increase in particles size with significant increase in drug encapsulation efficiency and loading. Increased PVA concentration resulted in reduced nanoparticle size, encapsulation efficiency and drug loading. Chitosan coated nanoparti cles were not significantly larger than uncoated particles, suggest ing only a thin coating layer. Reversal of zeta potential from negative for uncoated particles to positive for chitosan coated particles confirmed presence of chitosan coating. In vitro rel ease studies showed that TA -PLGA nanoparticles released at a similar rate to micronized drug particles. In vitro release of MF -PLA and CH MF produced a biphasic drug release profile. In vitro release determined that CH MF formulation released at the slowest rate, with only 43% drug release over 24 hours. Both chitosan coated (CH MF) and uncoated MF -PLA nanoparticles released drug at a slower rate compared to MF contained in the Asmanex Twisth alerTM.

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53 Chitosan coated particles, however, released at a much slower rate, only 20% burst release was observed, followed by slow drug release. Chitosan coated and uncoated nanoparticles prepared with the corticosteroid MF were able to release drug at a slower rate compared to MF contained in the Asmanex TwisthalerTM. Chitosan coated MF PLA nanoparticles were chosen for further development into nanocomposite microspheres due to slow release rates observed.

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54 Table 2 1 Advantages and disadvantages of various methods for production of nanoparticles Advantages Disadvantages Wet Milling Crystalline nanoparticles Contamination with grinding material Batch to batch variation Microbial growth @30C High pressure homogenization 40 500nm Small or large scale Aggregation and coalescence of nanosuspensions Changes in crystallinity Supercritical fluid process No organic solvent Depends on efficiency of atomization Hydrophobic:hydrophilic particles cannot be produced inadequate solvent systems Poorly water soluble drugs have poor CO2 solubility low particle production Resolved by high temperatures High pressures Solvent evaporation Spherical nanoparticles Monodispersed nanoparticles Ease of production Suitable for hydrophobic and hydrophilic drugs Use o f organic solvent Not suitable for industrial use

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55 Figure 21 Solvent evaporation technique to produce nanoparticles Dissolve polymer and drug in DCM PVA aqueous solution Sonicate at 60W for 5 mins O/W microemulsion Evaporation of DCM Centrifugation to collect nanoparticles Lyophilization

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56 Figure 22 Schematic detailing methods to coat nanoparticles with chitosan

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57 Figure 23 Determine encapsulation efficiency and drug loading of nanoparticles

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58 Figure 24 Influence of PLGA and [PVA] on encapsulation efficiency of TA into nanoparticles 0.0 10.0 20.0 30.0 40.0 50.0 60.0 70.0 80.0 90.0 100.0 200 300 400 % Encapsulation efficiencyPLGA (mg) 1% PVA 2% PVA 3% PVA

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59 Figure 25 Influence of PLGA and [PVA] on particle size of TA nanoparticles 0 50 100 150 200 250 200 300 400 Particle size (nm) PLGA (mg) 3% PVA 2% PVA 1% PVA

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60 Figure 26 SEM of TA -PLGA formulated with 200mg PLGA and 1% w/v PVA

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61 Figure 27 SEM of TA -PLGA formulated with 200mg PLGA and 2% w/v

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62 Figure 28 SEM of TA -PLGA formulated with 400mg PLGA and 1 % w/v PVA

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63 Figure 29 In vitro release TA -PLGA nanoparticles formulated with 300mg PLGA using 1% or 2% w/v PVA 0.00 20.00 40.00 60.00 80.00 100.00 0 100 200 300 400 500 600% TA releasedTime (mins) micronized TA 300mg PLGA 2% PVA 300mg PLGA 1%PVA

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64 Figure 21 0 In vitro release of TA -PLGA nanoparticles formulated with 400mg PLGA and 1% or 2% w/v PVA 0.00 20.00 40.00 60.00 80.00 100.00 120.00 0 100 200 300 400 500 600% TA releasedTime (mins) micronized TA 400mg PLGA 1%PVA 400mg PLGA 2%PVA

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65 Figure 21 1 SEM of MF -PLA nanoparticles incubated with chitosan 1% w/v

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66 Table 2 2 Particle size distr ibution for chitosan coated MF -PLA nanoparticles prepared with 0.1% or 1% w/v chitosan by either incubation or in situ coating with chitosan Method Particle size (nm) SD (nm) Incubation 0.1% w/v chitosan Bimodal Incubation 1% w/v chitosan Bimodal In situ 0.1% w/v chitosan 226 10 In situ 1% w/v chitosan Bimodal

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67 Table 2 3 Influ ence of chitosan coating to MF -PLA nanoparticles on encapsulation efficiency, drug loading, particle size and zeta potential Formulation Drug loading (%w/w) SD (%w/w) Particle size (nm) SD (nm) Zeta potential (mV) PLA nanoparticles 1.4 0.1 212 14 36.0 Chitosan coated PLA nanoparticles 2.0 0.1 226 10 13.7

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68 Figure 21 2 SEM of uncoated MF PLA nanoparticles formulated with 400mg PLA, 1% w/v PVA

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69 Figure 21 3 SEM of chitosan coated MF -PLA nanoparticles formulated with 400mg PLA, 1% w/v PVA and 0.1% w/v chitosan

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70 Figure 21 4 In vitro release of MF PLA, chitosan coated MF -PLA (CH -MF) and micronized MF 0 20 40 60 80 100 120 0 200 400 600 800 1000 1200 1400% MF releasedTime (mins) PLA -MF CH-MF Asmanex

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71 CHAPTER 3 DEVELOPMENT OF NANOC OMPOSITE MICROSPHERE S Introduction Although nanoparticles may be capable of provid ing increased pulmonary retention and slow drug release, they are not efficiently deposited within the airways. In order to deposit in the airways particles must have an MMAD in the respirable range of 1 -5 m [26] Nanoparticles must first be incorporated into micron sized carrier particles, upon exposure to the aqueous environment of the lung, these particles are released, and deposit nanoparticles onto the lung epithelium [102] Nanoparticles may be spray dried with an ex c ipient such as lactose to form nanocomposite microspheres with good flow properties [103-105] Ely et al developed effervescent spray dried nanoparticles, these contain an active method to breakdown microspheres once deposited in the lung; citric acid and carbonate react when delivered to the humid ai rways for form bubbles of carbon dioxide and disperse nanoparticles [106] Hadinoto et al investigated the influence of nanoparticle size, chemical nature and feed concentration of polystyrene and colloidal silica nanoparticles to be spray dried on the ability to form hollow nanocomposite particles and found that a particular nanoparticle concentration threshold must be reached in order to produce hollow microspheres composed of nanoparticles alone [107] Addition of phospholipids also reduced phagocytic uptake, nanocomposite particles have bee n formulated by spray drying nanoparticles with phospholipids however th e s e particles result ed in slower drug release after spray d rying [108] Spray dried gelatin nanoparticles were found to be significantly la rger following spray drying under conventional conditions in a commercially available instrument

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72 however no size change was seen for polybutylcyanoacrylate particles [105, 109] Spray dried PLGA dexamethasone nanoparticles were spray dried with 1,2Dipalmitoyl sn -Glycero 3 -Phosphocholine (DPPC) and hyaluronic acid. The resultant microspheres produced released dexamethasone at a slower rate compared to nanoparticles alone [103] Aspirin nanoparticles spray dried with increasing concentrations of phospholipids initially had similar burst release rates to nanoparticles alone, however, following the initial burst nanocomposite parti cles with higher phospholipid contents released at a slower rate [107] Similar findings have also been noted in other studies [103] Large porous lactose particles have been developed by the Edwards group, the advantage of these are reduced aggregation on storage compared to smaller nonporous particles with an equivalent MMAD [102] Lactose is a commonly used excipient for inhaled formulations, a number of studies have used lactose to form nanocomposite microspheres [105, 106, 109] PLGA nanoparticles spray dried with trehalose or lactose are able to release nanoparticles on contact with lung lining fluid. If these n anocomposite microspheres are produced by spray drying using inlet temperatures of 100 C or higher, it is no longer possible for nanoparticles to be released from the formulation [110] Commercially av ailable spray drying systems operate with inlet temperatures ranging from 100220C, however, this may be problematic for use with polymeric nanoparticles PLGA and PLA which are used to form nanoparticles ha ve glass transition temperatures of 45 50 C and 50 -60 C respectively thus spray drying at higher temperatures will result in changes in particle shape and possibly agglomeration of nanoparticles [111]

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73 Hypothesis We hypothesize that a spray dryer capable of operating under lower temperatures than commercially available will allow nanoparticles to be incorporated into microspheres without change s in particle morphology, size or in vitro release characteristics. During the drying of aerosol droplets within the spray dyer, nanoparticles are protected from excessive heat due to evaporation of water. Onc e the aerosol is dried, nanoparticles may be exposed to high temperatures resulting in agglomeration of particles as well as changes in drug release rates, thus i t is especially important that the outlet temperature of the spray dryer should be lower than the Tg of the polymer nanoparticles. Lactose will be used as an excipient; we hypothesize that nanoparticles will be freely re -dispersed from nanocomposite microspheres composed of lactose. Materials and Methods Chemicals MF was donated by Ipca laboratories ltd (Mumbai, India). PLGA and PLA were purchased from Lactel Absorbable Polymers (Pelham, AL, USA). Extra -fine lactose was donated by EM industry (Hawthorne, NY, USA). DCM and ACN were purchased from Fisher Scientific (Pittsburgh, PA, USA). PVA and medium molecular weight chitosan were obtained from Sigma Chemical Co. (St. Louis, MO). Development of Spray Dryer and O ptimization of O perating C onditions DL -PLA has a glass transition temperature (Tg) of 50-60 C, a novel spray dryer capable of operating at temperatures lower than the Tg of the polymers was developed (Figure 3 -1) ; this allowed nanoparticles to be spray dried without causing particles to aggregate during the process. Operating conditions of the spray dryer will be optimized to al low

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74 dried microspheres to be collected using an Anderson cascade impactor (Copley Scientific UK ) Initial ly, PLA -TA nanoparticles were spray dried with lactose to determine the lowest temperatures and additional parameters under which dried microspheres may be produced. Parameters such as sweep air flow, starting temperature of the drying chamber, nebulizer air flow, nebulizer flow rate were varied. Particle morphology was observed by SEM to determine particle morphology and incorporation of nanoparticl es into microspheres. During method development in order to determine operating parameters for spray drying, t otal collection of TA -PLA nanoparticles on the ACI w as evaluated by weight of particles deposited on each stage. In addition, each stage w as rin sed with 5 mL MeOH and total TA content will be evaluated following dilution with DDW to a produce 50:50 MeOH:DDW solution Total mass of nanoparticles collected on each stage was determined This allowed a fast screening method to determine the composition of nanocomposite microspheres produced as well as an approximation of the MMAD of the formulation. The spray dryer design is shown in Figure 3 -1. Warm dry air is directed into the heating chamber from the bottom of the instrument. The port pl aced at the bottom of the heating chamber is angled to allow the warmed dry sweep air to produce a vortex through the heating chamber; preventing nebulized aerosol droplets impinging the spray dryer walls. This results in reduced loss of formulation withi n the spray dryer, to improve the final yield of spray dried nanoparticles. In addition to this, it increases the time the aerosol droplets reside in the heating chamber, allowing a more gentle heat to be applied. In addition width and height of the aerosol generated using the nebulizer

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75 was measured, as a result the diameter of the drying chamber was designed to be wider than the aerosol plume created by the glass nebulizer to reduce impaction and loss to the chamber walls. Inlet temperature wa s maintained in the range 7580 C while outlet temperature wa s less than 35 C. Parameters such as temperature, nebulizer air flow, flow rate of nanoparticle suspension to be aerosolized and air flow through the spray drying device were optimized in order to allow complete drying of the microspheres without subjecting nanoparticles to temperatures above Tg. Optimal conditions were found to be with inlet airflow of 15 L/min, 0.5 ml/min flow of nebulizer suspension with the whole system under slightly negative pressure Inlet temperatures less than 75 C resulted in incomplete drying of the microspheres, which could be observed by the collection of wet particles in the ACI. Higher flow rates of the nebulized suspension also resulted in wet particles being produced. Th e nanoparticle/lactose suspension is sprayed into a heated drying chamber using a type A concentric circle glass nebulizer (Meinhard, Golden, CO) at a rate of 0.5 mL/min under 20 psi with a corresponding air flow of 0.8 L/min, shown in Figure 32. The neb ulizer is able to operate at a maximal pressure of 30 psi, with a linear correlation between air flow rate and pressure between 15 and 80 psi, (Air flow (L/min) = 0.0328xPressure(psi) + 0.1401 with R2 = 0.9997). The heated chamber was maintained under a s lightly negative pressure in order to improve aerosol drying. Warm dry sweep air at a flow rate of 15 L/min was used to create a vortex within the drying chamber to prevent aerosol droplets from impinging onto the chamber walls. Inlet and outlet temperatures were approximately 78 C and <35 C respectively; outlet temperatures should be maintained below the Tg of the polymer in order to prevent

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76 nanoparticle agglomeration. Typically commercial spray dryers operate at much higher temperature, for example lact ose is spray dried at an inlet temperature of 160 C and an outlet temperature of 105 C this method would be unsuitable for use with polymeric nanoparticles [112] An Anderson cascade impactor was set up inline with the spray dryer as a method of both collecting the spray dried formulation and also to selectively collect the respirable fraction. Air flow through the impactor was set to 20 L/min, following spray drying, the formulation was retrieved from stages 27. Particles collected are expected to have a MMAD below 6. 9m. Spray Dried Chitosan Coated MF PLA Nanopar ticles Optimal spray dryer conditions determined were used for further development of spray dried nanocomposite microspheres. In Chapter 2, chitosan coated PLA nanoparticles were shown to produce the slowest drug release compared to uncoated nanoparticles or MF contained in the Asmanex formulation (Figure 2 -1 4 ), as a result were spray dried with lactose. Chitosan coated nanoparticles (CH MF) were suspended in a lactose solution, then spray dried. Total soli d feed content to be nebulized w as initially set to 5% w/v while the CHMF:lactose composition was varied (10:90, 25:75, 50:50 75: 25 and 90:10). MMAD, particle morphology, nanoparticle entrapme nt efficiency and in vitro release characteristics of spray dried CH -MF nanoparticles were determined as described below Nanoparticle Entrapment Efficiency and Loading into Nanocomposite Microspheres In order to determine the percentage nanoparticle content of the nanocomposite microspheres as well as th e nanoparticle encapsulation efficiency 5 mg of the nanocomposite microspheres were weighed and dissolved in DCM on the orbital shaker

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77 for overnight DCM was placed in the vacuum centrifuge (Jouan RC10.10) to dryness the dried residue was dissolved in 1 mL ACN:DDW 60:40 and analyzed by HPLC (Hewlett Packard Series 1050). Drug loading and encapsulation efficiency were determined by the equations below. ( % / ) = ( ) ( ) 100 = ( % / ) ( % / ) 100 Morphology of Nanocomposite Microparticles SEM was performed to observe particle morphology using Scanning Electron M icroscope (SEM) JEOL JSM -6335F instrument (Major Analytical Instrument Center (MAIC), UF, Gainesville, FL). The spray dried formulation was placed on a carbon stub which was coated with carbon using a vacuum evaporator, SEM was performe d at 2 kV. Lactose alone was spray dried under the same conditions to allow comparison of particle morphology of spray dried particles which do not contain any nanoparticles. Determine MMAD of Nanocomposite Microparticles Particle MMAD of spray dried hybr id particles was determined using an 8-stage, nonviable Anderson cascade impactor (Copley Scientific), stainless steel collection plates were used, the inhaler device was coupled to the impactor using a tailor made adapter. The operational airflow used was 39 L/min for 6 seconds per actuation. Surfaces of the particle collection sites on the ACI were coated with an ethanolic solution of brij 35 in glycerol to avoid bias caused by particle bounce [113] Each of the plates was coated with 50 L br ij 35 solution, with the pre-seperator being coated with 100 L to reduce the bounce effect ethanol will evaporate before HPLC analysis The sample induction

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78 port (SIP), high top, pre-separator and plates from each stage were washed with 5 mL acetic acid buffer pH 4, to remove the chitosan coating, this was followed by incubation with 5 mL acetonitrile to release MF from within the PLA matrix. The samples were centrifuged then analyzed by HPLC as described above. Two capsules were half filled with 10 mg of the spray dried formulation; the capsule was placed in the center chamber of the device with the overlapping end pointing upwards and delivered via the HandiHaler (Figure 3 4 ). Capsules were pieced in the HandiHaler however the size of the holes produced w as designed to be smaller diameter than that used in the commercial device, in order to prolong the residence/delivery time of the formulation within the capsules and thereby allowing for a longer deagglomeration process of the formulation parti cles. The respirable fraction of the spray dried nanoparticles formulation was compared with that of the commercially available MF formulation delivered from the Asmanex inhaler device (Figure 3 -5) The impactor will not give a continuous distribution of particles, but will categorize particle size into certain size ranges [114] The respirable dose was calculated based on particles with MMAD <4 .9 m by ACI analysis, based on MF collection from stage 2 and below, based on the particle size cutoff shown in Table 3 1 The Metered Dose (MD) was based on the MF content using the actual amount of formulation weighed into the capsules for inhalation. The Delivered Dose (DD) was determined based on the total recovered MF from all stages of the ACI, pre -separator Sample Induction Port (SIP), high top and adapter. Drug remaining in the device and capsules were not included within the DD. Fine particle dose (FPD) was determined by the summation of all MF collected from stage 2 and below and was compared to DD to determine FPF.

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79 = ( ) ( ) 100 In vitro drug Release S tudy In vitro release was performed in 100 mL 1% v/v tween 80 in PBS, shaken at 30 rpm in a hot shaker (Bellco biotechnology, Vineland, NJ) at 37 C over 24 hours under constant sink conditions as previously described in Chapter 2 Tween 80 increase d the saturation concentration of MF in the dissolution media to 25 g/ml at 37 C. In order to maintain sink conditions maximal concentrations obtained did not exceed 10% of the saturation concentration. At specific time points (0, 15, 30, 90, 12 0, 180, 240, 360, 600 and 1440 minutes), 1 mL was removed using a pipette and subsequently filtered using a 0.02 m filter (Whatman Anotop plus filter), 1 mL fresh buffer was replaced into the dissolution media. HPLC was used to quantify released drug concentrations using a calibration curve ranging from 0.5-10 ug/mL with an R2>0.996. Spray dried CH MF formulations were compared to MF from the Asmanex TwisthalerTM. Results and D iscussion Development of Spray Dryer and O ptimization of Operating C onditions A spray dryer capable of operating with inlet temperatures less than the Tg of PLGA or PLA was designed. Additionally it was able to spray dry formulations using mg quantities of materials. Initially PLA -TA nanoparticles were used to optimize the spray dryer operational parameters, SEM of this formulation can be seen in Figure 3-9 Figure 36 shows MMAD of particles collected directly from the ACI, a large fraction of the total spray dried microspheres obtained are less than 5.8 m, Figure 3 7 also shows this as the total mass collected at each stage. Figure 38 shows that in general the total nanoparticle:lactose content did not vary between the collection stages, with an

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80 exception of a much lower incorporation of nanoparticles into microsphe res <0.65 m. This may be due to exclusion of nanoparticles from microspheres that are very small, as the nanoparticles are also sized at least 200 nm. Initially mass of formulation deposited on the collection surfaces were determined by gravimetric anal ysis as a convenient method to quantify larger amounts of collected formulation. However, for smaller masses, it becomes more accurate to perform analysis by HPLC as gravimetric analysis is not able to differentiate between formulation weight and water weight therefore spray dried CH -MF nanoparticles were further analyzed by HPLC [115] Spray Dried Chitosan Coated MF PLA Nanoparticles Influence of CH -MF:lactose Ratio on M orphology of S pray D ried M icrospheres Chitosan coated MF PLA (CH MF) nanoparticles were used to optimize composition of nanocomposite microspheres T he effect of increasing nanoparticle:lactose ratio on particle morphology was investigated. CH-MF n anoparticle content was varied at 10%, 25%, 50% and 75% of the total feed of 5 % w/v being spray dried. Higher nanoparticle content was not possible as a result of blockage of the nebulizer nozzle with the use of 90% nanoparticle composition. At low composition of 10% nanoparticles, m icrospheres formed were spherical as shown in F igure 3-10 however, as nanoparticle content increased, particles appeared more deflated and hollow as shown in F igures 3 11, 3 12 and 313. Advantages of spray dried hybrid particles containing higher nanopa rticle content include a lower mass to be inhaled as well as production of hollow particles which in turn will have lower MMAD as in comparison to solid particles with similar particle size. Hollow particles tend to have larger particle sizes compared to non -porous particles of the same MMAD. This property allows deposition of these particles in the same regions of the airways,

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81 however less aggregation should be expected on storage. As a result of this, chitosan coated nanoparticles w ere developed using a high nanoparticle:lactose ratio of 75:25. The high solid feed content of 5% w/v tended to lead to clogging of the nebulizer nozzle, therefore particles were spray dried with a solid feed concentration of 1.25% w/v. Lactose alone was spray dried at a concentration of 1.25% w/v to determine surface morphology of these microspheres in the absence of incorporated nanoparticles. Figure 3 -14 shows the SEM of spray dried lactose alone, microspheres have a smooth surface in comparison to lactose spray dried wi th nanoparticles. The resultant microspheres were still produced using a ratio 75% CH MF nanoparticles and 25% lactose, therefore the final formulation will have the same nanoparticle and drug loading. Spray Dried Chitosan Coated MF PLA Nanoparticles Nanoparticle Incorporation Chitosan coated nanoparticles were spray dried with lactose 7 5% and 2 5% r espectively, with a solid feed concentration of 1.25% w/v. Solid feed content was reduced from 5% w/v from previous studies with uncoated MF PLA nanopartic les to 1.25% w/v as chitosan coated nanoparticles tended to clog the nebulizer at high co ncentrations. A relatively high nanoparticle:lactose ratio was chosen in order to produce a formulation which contained a clinically appropriate dose and could be a dministered by one or two inhalations after filling capsules for use in the HandiHaler. Theoretical MF loading into the spray dried CH MF formulation is 1.5% w/w, actual MF loading was very similar at 1.43% w/w (SD=0.08% w/w). It was determined that 9 6 % (SD=6%) of the total nanoparticles used were incorporated into the final spray dried formulation SEM of the formulation contained within the Asmanex TwisthalerTM is s hown in Figure 3 15, this shows a blend of lactose with MF.

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82 Spray Dried Chitosan Coated MF PLA Nanoparticles MMAD MMAD of spray dried CH MF nanoparticles were characterized by ACI to determine FPF of the delivered dose from the HandiHaler at 39 L/min for 6 seconds The United States Pharmacopeia recommends that DPIs should be test ed at a flow rate which corresponds to a pressure drop across the inhaler of 4 kPa [116] As the Handi Haler produces a pressure drop of 4 kPa at 39 L/min, this was the air flow tested On inspiration the capsule vibrates inside the center chamber, this results in mechanical agitation of powder aggregates contained and results in their break up [117] Increasing retention of the formulation within the capsule before delivery to the lung allows increased time over which deaggregation may occur. In order to increase the time tha t the spray dried formulation is mechanically agitated within the device; holes used to pierce the capsule were customized to be smaller than those usually made with the HandiHaler, while allowing the capsule to empty with one inhalation. FPF of MF conta ined in Asmanex TwisthalerTM was used as a control for the spray dried CH -MF nanoparticles. The Metered Dose (MD) and Delivered Dose (DD) of MF delivered from the Asmanex TwisthalerTM produced by ScheringPlough is available as 220 g or 440 g and 200 g or 400 g respectively [118] The HandiHaler is a high efficiency inhaler and was used as it is capable of delivering large doses up to 50 mg by inhalation [1 19] This allow s a clinical dose of MF from the spray dried chitosan coated nanoparticle formulation to be administered with one or two inhalations. In addition, prepared spray dried formulations can be packaged into capsules with relative ease in com parison to other DPI devices such as the FlixotideTM DiskhalerTM which would require special blister packing equipment to be used [120]

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83 FPF on DD from the spray dried chitosan coated nanoparticle formulation was 14.6% (SD=4.3%), whereas the Asmanex TwisthalerTM had 20.0% (SD=4.7%). Three batches of spray dried formulation were compared in order to determine batch to batch variability of the spray drying process (n=3 per batch), shown in Figure 3 16 MMAD of MF from the TwisthalerTM was evaluated in order to make comparisons with the FPF on DD of the spray dried nanoparticle formulation with a known commercial formulation. Literature values for FPF of the Asmanex TwisthalerTM are reported to be 39.9% (SD=2.5) and 35.6% (SD=3.4) at 200 g and 400 g DD with an airflow of 60 L/min. The fine particle dose was however defined as particle with MMAD <6.5 m [121] Differences between ACI conditions will account for the lower FPF observed with the Asmanex compared to the literature reported values. To summarize, FPF of the DD from the spray dried CH MF formulation is similar to commercially available MF from Schering -Ploughs Asmanex TwisthalerTM. Spray Dried Chitosan Coated MF PLA Nan oparticles In vitro release MF blended with lactose was removed from the Asmanex TwisthalerTM, Figure 3 15. At 1 hour 100% of MF from the Asmanex is released, in comparison, spray dried CH-MF exhibits an initial burst, followed by slower release. Spr ay dried CH -MF shows a slightly greater burst effect compared to CH -MF nanoparticles alone of 30% and 20% respectively. At 24 hours approximately 70% of the total MF contained within the spray dried formulation is released, however, CH MF nanoparticles alone only released 43 % of the total MF content at this time point (Figure 317) Lyophilized n anoparticles must be re dispersed in a lactose solution before they can be spray dried. It is possible that during this process, MF is able to disperse to the surface of the CH MF nanoparticles, leading to a greater burst release. Following the initial phase, release rate from CH -MF

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84 nanopart icles and s pray dried formulation were similar, indicating that nanoparticles did not aggregate during the spray drying process. In vitro release of CH -MF nanoparticles did no t change following spray drying. Additionally, the spray dried formulation released MF at a much slower rate compared to the commercial MF prepared in the Asmanex TwisthalerTM. Studies conducted by other groups observed slower drug release from PLGA nanoparticles spray dried with lipids, however it is unclear if this is a result of high temperatures used or a property of the lipids used [102, 107] Conclusion The spray dryer was designed to operate at outlet temperatures below the Tg of PLA or PLGA. Spray dried particle morphology was influenced by nanoparticle:lactose composition used. Spherical microspheres with MMAD in the respirable range were produced containing nanoparticles. Increase in nanoparticle:lactose ratio resulted in ch anges in microsphere morphology. Lower nanoparticle:lactose ratios produced spherical microspheres whereas higher nanoparticle concentrations produced hollow particles as shown by SEM Spray dried chitosan coated MF -PLA nanoparticles composed of 75% nanoparticles and 25% lactose are able to provide a clinically appropriate dose in approximately two inhalations using 20 mg of the formulation from the HandiHaler FPF on DD of the spray dried chitosan coated MF -PLA nanoparticles was 14.6% (SD=4.3%), whereas the Asmanex TwisthalerTM was 20.0% (SD=4.7%). In vitro release of MF from the spray dried chitosan coated MF -PLA nanoparticles was much slower than from the formulation contained within the Asmanex TwisthalerTM. A biphasic release profile was observed for spray dried CH MF hybrid particles

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85 After CH MF nanoparticles were spray dried, initial burst release observed was faster than for CH MF nanoparticles alone. Following the initial burst however, the release rate s were found to be similar.

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86 Figure 31 Spray dryer, designed to operate under lower temperatures than commercially available instruments

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87 Figure 32 TR -30A3 nebulizer, type A flush capillary lapped nozzle 30 psi, 3 mL/min [122]

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88 Figure 33 Schematic to determine incorporation efficiency of nanoparticles into spray dried formulation Lactose MF loading into CH MF nanoparticles (2% w/w) Spray dried: 25:75 lactose:CH MF nanoparticles Theoretical MF loading (1.5% w/w) Incorporation Efficiency of nanoparticles = (Actual/Theoretical drug loading)*100

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89 Figure 34 HandiHaler (1) dust cap, (2) mouthpiece, (3) mouth piece ridge, (4) base, (5) green piercing button, (6) center chamber, (7) air intake vents [123]

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90 Figure 35 Cross section of Asmanex TwisthalerTM [121]

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91 Table 3 1 MMAD cutoff for ACI a nalysis based on air flow rate (L/min) L/min 20.0 22.5 28.3 30.0 39.0 45.4 60.0 Stage 0 17.810.7 16.810.1 15.09.0 14.68.7 12.87.7 11.87.1 10.36 1 10.7 6.9 10.1 6.5 9.0 5.8 8.7 5.6 7.7 4.9 7.1 4.6 6.2 4.0 2 6.9 5.6 6.5 5.3 5.8 4.7 5.6 4.6 4.9 4.0 4.6 3.7 4.0 3.0 3 5.6 3.9 5.3 3.7 4.7 3.3 4.6 3.2 4.0 2.8 3.7 2.6 3.0 2.2 4 3.9 2.5 3.7 2.4 3.3 2.1 3.2 2.0 2.8 1.8 2.6 1.7 2.3 1.0 5 2.5 1.3 2.4 1.2 2.1 1.1 2.0 1.1 1.8 0.9 1.7 0.9 1.4 0 6 1.3 0.8 1.2 0.8 1.1 0.7 1.1 0.7 0.9 0.6 0.9 0.6 0.8 0 7 0.8 0.5 0.8 0.4 0.7 0.4 0.7 0.4 0.6 0.3 0.6 0.3 0.5 0 Filter <0.5 <0.4 <0.4 <0.4 <0.3 <0.3 <0.3

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92 Figure 36 MMAD of nanocomposite PLA -TA particles 0 5 10 15 20 25 30 5.8 9 4.7 5.8 3.3 4.7 2.1 3.3 1.1 2.1 0.65 1.1 0.43 0.65 % nanocomposite particlesMMAD (um)

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93 Figure 37 Total PLA -TA nanoparticles for each MMAD range 0 2 4 6 8 10 12 5.8 9 4.7 5.8 3.3 4.7 2.1 3.3 1.1 2.1 0.65 1.1 0.43 0.65 Formulation mass (mg) MMAD (m)

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94 Figure 38 Percentage of PLA -TA nanoparticles at each particle cutoff compared with total weight 0.00 1.00 2.00 3.00 4.00 5.00 6.00 7.00 8.00 9.00 10.00 5.8-9 4.7-5.8 3.3-4.7 2.1-3.3 1.1-2.1 0.65-1.1 0.43-0.65% Nanoparticles of total nanocomposite particle formulationMMAD (um)

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95 Figure 39 SEM of PLA -TA nanoparticles spray dried with lactose

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96 Figure 310 Spray dried chitosan coated MF -PLA nanoparticles containing 10% of total solid feed as nanoparticles

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97 Figure 311 Spray dried chitosan coated MF -PLA nanoparticles containing 25% of total solid feed as nanoparticles

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98 Figure 31 2 Spray dried chitosan coated MF -PLA nanoparticles containing 50% of total solid feed as nanoparticles

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99 Figure 31 3 Spray dried chitosan coated MF -PLA nanoparticles containing 75% of total solid feed as nanoparticles

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100 Figure 31 4 Spray dried lactose, 1.25% w/v solid feed content

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101 Figure 31 5 SEM of formulation contained in Asmanex TwisthalerTM

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102 Figure 31 6 FPF of DD (%) of spray dried chitosan coated nanoparticles, to compare batch to batch variability (n=3) 0 2 4 6 8 10 12 14 16 18 20 22 24 26 28 30 Batch 1 Batch 2 Batch 3 FPF of DD (%)

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103 Figure 31 7 In vitro release of MF from spray dried CH MF nanoparticles and Asmanex 0 20 40 60 80 100 120 0 200 400 600 800 1000 1200 1400MF released (%)Time (mins) Spray dried CH MF nanoparticles Asmanex MF CH MF

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104 CHAPTER 4 SUMMARY Polymeric nanoparticles containing PLGA and the model drug TA were developed using t he solvent evaporation. Spherical unimodal TA -PLGA nanoparticles were produced using this technique. It was shown that various process parameters could influence particle size, drug encapsulation an d drug release. Increasing p olymer content was shown to increase encapsulation of TA in TA -PLGA nanoparticles and result in a trend of increasing particle size. As PVA surfactant concentration was increased, encapsulation efficiency decreased and smaller nanoparticles were formed. These TA -PLGA nanoparticles were able to release at a similar rate as micronized TA. Further developments were made using PLA to form nanoparticles. Mometasone furoate was used as this is a more lipophilic drug and would show slower drug release from pol y mer nanoparticles. Drug release was further reduced by coating MF -PLA nanoparticles with chitosan. MF released at a substantially slower rate from chitosan coated nanoparticles compared with MF -PLA nanoparticles and micronized MF that was used as a control. Nanocomposite particles were developed using a spray dryer that was developed; this spray dryer operated with outlet temperatures below the Tg of PLA or PLGA making it suitable for spray drying polymer nanoparticles. The spray dryer was able to produce particles with MMAD in the respirable range. Nanocomposite microspheres were able to deliver a clinically relevant dose with two SPIRIVA capsules filled with 10 mg of the formulation using the HandiHale r device. SPIRIVA capsules were placed in the center chamber of the HandiHaler and pierced by pressing the button to a pre determined level. Holes size produced in the capsule was customized to increase

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105 retention within the device for a longer time period and therefore increase the time over which deaggregation may occur. As a result of both the formulation and optimizing the inhaler device for use with this formulation, the FPF was similar to that from the Asmanex TwisthalerTM. MF release from the spray dried CH MF was compared to the nanoparticles before spray drying and also MF formulation contained within the Asmanex TwisthalerTM. MF was release rapidly from the Asmanex with 100% release by 1 hour. In comparison both CH MF and the spray drie d formulation exhibited a biphasic release profile. CH MF nanoparticles showed a burst release of approximately 20% compared to 30 % from the spray dried formulation. A possible explanation for this may be as a result of having to re -disperse lyophilized nanoparticles in a lactose solution before spray drying. This could cause some of the MF to diffuse to closer to the surface of the nanoparticles and result in a greater burst effect observed with spray dried nanoparticles. Following the initial burst, b oth the CH MF nanoparticles and spray dried formulation release with a similar rate. In vitro release from nanoparticles was not altered by the spray drying process.

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106 LIST OF REFERENCES 1. Summary Health Statistics for U.S. Adults: National Health Interview Survey, 2008. US Department of Health and Human Services, Centers for Disease Control and Prevention, National Center for Health Statistics, 2009. 2. Summary Health Statistics for U.S. Children: National Health Interview Survey, 2008. US Department of Health and Human Services, Centers for Disease Control and Prevention, National Center for Health Statistics, 2009. 3. Eder, W., Ege, M.J., and Mutius, E.V., The Asthma Epidemic. N Engl J Med, 2006. 355(21): p. 22262235. 4. Global strategy for asthma management and prevention. Global Initiative for Asthma, 2008. 5. Elias, J.A., Airway Remodeling in Asthma Unanswered Questions. Am. J. Respir. Crit. Care Med., 2000. 161 (3): p. S168-171. 6. Corrigan, C.J., et al., The ADMIT series Issues in Inhalation Therapy. 3) Mild persistent asthma: the case for inhaled corticosteroid therapy. Primary Care Respiratory Journal, 2009. 18(3): p. 148158. 7. Kelly, H.W., Potential adverse effects of the inhaled corticosteroids. Current reviews o f allergy and clinical immunology, 2003. 112 : p. 10. 8. Weldon, D., The effects of corticosteroids on bone growth and bone density. Annals of Allergy, Asthma and Immunology, 2009. 103: p. 3 11. 9. Hubbard, R. and Tattersfield, A., Inhaled Corticosteroids, Bone Mineral Density and Fracture in Older People. Drugs & Aging, 2004. 21 (10): p. 631638. 10. Peters, S.P., Safety of Inhaled Corticosteroids in the Treatment of Persistent Asthma. Journal of the National Medical Association, 2006. 98(6 ): p. 11. 11. Roland, N.J., Bhalla, R.K., and Earis, J., The Local Side Eff ects of Inhaled Corticosteroids Chest, 2004. 126(1): p. 213 -219. 12. Hochhaus, G., New Developments in Corticosteroids. Proc Am Thorac Soc, 2004. 1 (3): p. 269274. 13. Thorsson, L., Edsbacker, S., and Conradson, T.B., Lung deposition of budesonide from Turbuhaler is twice that from a pressurized metered -dose inhaler P -MDI. Eur Respir J, 1994. 7 (10): p. 18391844. 14. Derendorf, H., et al., Relevance of pharmacokinetics and pharmacodynamics of inhaled corticosteroids to asthma. Eur Respir J, 2006. 28(5): p. 10421050.

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117 BIOGRAPHICAL SKETCH Gina Patel was born in 1981 London, England. She graduated from the Universit y of Bath, Bath, UK in 2003 as a M aster of Pharmacy. During her studies she participated in an internship at the University of Florida under the supervision of Dr Hochhaus. After qualification as a pharm acist she worked as a hospital pharmacist at Musgrove Park Hospital, Taunton, UK before joining the Department of Pharmaceutics, University of Florida.