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Review
10.1517/17425240802141568 © 2008 Informa UK Ltd ISSN 1742-5247 1
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Biodegradablepolymeric
nanocarriersforpulmonary
drugdelivery
Erik Rytting, Juliane Nguyen, Xiaoying Wang & Thomas Kissel†
Philipps-Universität Marburg, Institut für Pharmazeutische Technologie & Biopharmazie,
Ketzerbach 63, D-35032 Marburg, Germany
Background: Pulmonary drug delivery is attractive for both local and
systemic drug delivery as a non-invasive route that provides a large surface
area, thin epithelial barrier, high blood flow and the avoidance of first-pass
metabolism. Objective: Nanoparticles can be designed to have several
advantages for controlled and targeted drug delivery, including controlled
deposition, sustained release, reduced dosing frequency, as well as an
appropriate size for avoiding alveolar macrophage clearance or promoting
transepithelial transport. Methods: This review focuses on the development
and appli cation of biodegradable polymers to nanocarrier-based strategies
for the delivery of drugs, peptides, proteins, genes, siRNA and vaccines by
the pulmonary route. Results/conclusion: The selection of natural or
synthetic materials is important in designing particles or nanoparticle
clusters with the desired characteristics, such as biocompatibility, size,
charge, drug release and polymer degradation rate.
Keywords: biodegradable polymers, gene and siRNA delivery, nanoparticles, peptides,
proteins, pulmonary drug delivery
Expert Opin. Drug Deliv. (2008) 5(6):1-11
1. Introduction
Pulmonary drug delivery is attractive for several reasons. It is the obvious choice
for the local administration of drugs to treat disease locally within the lung, but
there are several advantages of employing the pulmonary route to achieve systemic
delivery of therapeutics.
Diseases that can be targeted with local pulmonary administration include
chronic obstructive pulmonary disease (COPD), asthma, cystic fibrosis, infectious
diseases, tuberculosis and lung cancer [1]. Delivering drugs via the lungs
also provides a non-invasive route of delivery for targeting the systemic circulation,
as the lungs provide a large surface area, a thin epithelial barrier, high blood
flow, and less enzymatic activity compared to other areas in the body [2-4].
First-pass metabolism can be avoided by pulmonary administration, which can be
especially useful for biopharmaceuticals, which are often extensively degraded
following oral delivery [4,5].
1.1 Pulmonaryphysiology
After inhalation through the nose or mouth, air first enters the trachea, which
divides into two main bronchi. Bifurcations continue through 23 stages before
the alveolar sacs are reached. Several changes in the cellular milieu occur when
moving from the bronchial regions to the alveolar regions deep in the lung,
including differences in epithelial cell type, airway thickness and lining fluid [6].
The airway epithelium contains ciliated cells, secretory cells and basal cells; the
alveolar epithelium is mainly comprised of Type I and Type II alveolar cells.
1. Introduction
2. Biodegradable
nanocarrier materials
3. Pulmonary delivery of
active agents
4. Conclusions
5. Expert opinion
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Narrower airways result in narrower diffusion distances. The
total distance spanning the air–blood barrier decreases from
about 30 – 40 μm in the bronchial region to approximately
500 nm in the alveolar region [4,6]. The mucus layer lining
the upper airways is about 5 – 10 μm thick, whereas the
lung surfactant lining fluid secreted by the Type II alveolar
cells has a layer thickness of only 50 – 80 nm [6].
Barriers to particle deposition include the physical
defense of the oropharyngeal region and the bronchial tree.
Other barriers to particle and drug transport in the airways
include the mucus layer, the alveolar lining fluid, epithelial
cells, basement membrane, the mucociliary escalator,
macrophage clearance and proteolytic degradation [4,5]. Tight
junctions in the distal airways are not as tight as the tight
junctions in the bronchial and alveolar epithelia, suggesting that
targeting these distal airways may result in higher bioavail-
abilities in the systemic circulation [7]. Alveolobronchial
clearance is slower in the peripheral regions, and macrophage
clearance is reported to be minimal for particles less
than 260 nm [3,5]. Metabolism in the lung occurs by
peptidases and Phase I enzymes; these metabolizing enzymes
are found mostly in Clara cells, alveolar Type II cells
and alveolar macrophages [3,6].
1.2 Nanocarriersforpulmonarydelivery
Nanoparticles have gained increasing attention for pulmonary
drug delivery, due to their advantages for targeted deposition,
bioadhesion, sustained release and reduced dosing frequency
for improving convenience to the patient [2]. Some incentives
for using nanoparticles for the controlled delivery of drugs,
peptides, proteins, genes, siRNA and vaccines in the lung
include having an appropriate size for avoiding alveolar
macrophage clearance and promoting transepithelial transport.
Nanocarriers used for pulmonary applications also include
liposomes, solid lipid nanoparticles and nanotubes, but since
this review is limited to polymeric nanocarriers, the reader is
referred to other excellent manuscripts of interest [8-10]. Lung
targeting following intravenous nanoparticle administration
also falls outside the scope of this review, but this has been
discussed previously [11]. Due to their small size, most
nanoparticles would be exhaled, but these multifunctional
particles can be delivered to the lung by nebulization or by
the incorporation of the nanoparticles into larger particles
with an appropriate aerodynamic diameter by flocculation [12],
spray drying, or other means.
Regional deposition of particles delivered to the lung
depends on several factors, including particle properties such
as aerodynamic diameter, charge, surface properties and
hygroscopicity, as well as temperature, breathing pattern and
the timing of the aerosol pulse injection within the breathing
cycle [4]. The aerodynamic diameter (dae) is a function
of size, shape and density. Porous particles will have a
smaller dae than their physical diameter would suggest, and
for non-spherical particles the dae is mostly dependent upon
the short axis and the magnitude of the aspect ratio [5].
Shape does not only affect deposition by its influence
on the dae, but other factors too, as particle fibers are not
as easily cleared by alveolar macrophages compared to
spherical particles [13].
The aerodynamic diameter (dae) has an important
influence on particle destination. The optimal size for
deposition in the deep lung for systemic delivery is
approximately 1 – 3 μm [5]. Particles larger than 5 – 10 μm
result in oropharyngeal deposition, and are more likely to
be swallowed than to reach the lung. Particles smaller
than 1 μm will likely be exhaled. For particles between
1 and 5 μm, the smaller particles generally reach the deeper
parts of the lung, and the larger particles land in the upper
airways [3]. Particles around 1 – 2 μm have a higher chance
of crossing the air–blood barrier, and particles smaller
than 150 nm encounter delayed lung clearance, increased
protein interactions and more transepithelial transport
compared to larger particles [13]. Particle size may also
affect particle degradation and drug release rates.
Surface charge is another important property to consider
in particle design. Low surface energy is needed to
avoid particle agglomeration [5,13]. Electrostatic interactions
are also possible between the alveolar wall and oppositely
charged particles, but this depends on hydrophobicity
and humidity [13].
2. Biodegradablenanocarriermaterials
The design and synthesis of biodegradable polymeric
materials that will provide the appropriate nanocarrier chara-
cteristics for temporal and spatial distribution of drug in
the lung has been pursued extensively. Nanocarrier targeting
to the lung tissue based on particle size and surface charge
is an important aspect for material selection and design,
but the release of the active agent will also depend on
its distribution in the nanocarriers and the degradation
rate of the polymer.
A number of synthetic and natural polymers have been
utilized in formulating biodegradable nanoparticles [14].
Synthetic polymers have the advantage of sustaining the
release of the encapsulated therapeutic agent over a period
of days to several weeks. Natural polymers have a compara-
tively short duration of drug release. Polymers used for the
formulation of nanoparticles include natural polymers such
as albumin, gelatin, alginate, collagen, cyclodextrin and chitosan;
synthetic polymers used for pulmonary applications include
poly(lactide-co-glycolide) (PLGA) copolymers, polyacrylates
and polyanhydrides. These polymers, together with novel
polymers representing modifications to PLGA, are outlined
below. The applications of these polymers to pulmonary
drug delivery are then described in the following section.
2.1 PLAandPLGA
Polylactides (PLA) and poly (D,L-lactide-co-glycolide)
(PLGA) have been the most extensively investigated for drug
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delivery [14]. As polyesters in nature, these polymers undergo
hydrolysis upon implantation into the body, forming
biologically compatible moieties (lactic acid and glycolic
acid) that are removed from the body by the citric acid
cycle. The degradation products are formed slowly and do
not affect normal cell function. These safe, non-toxic
polymers are currently being widely used in drug delivery
systems and tissue engineering research.
More details regarding the biocompatibility and bio-
degradation of PLA and PLGA in drug delivery applications
are presented in a review by Anderson and Shive [15]. The
release of drug in the PLGA matrix is controlled by diffusion
of the drug through the matrix and by degradation of
the polymer [15]. PLGA’s degradation rate is affected by the
copolymer composition and the molecular weight, and the
release of drug can thus be varied from weeks to months [16].
Several methods such as solvent evaporation, nanoprecipitation
and multiple emulsions allow formulations of small particles
on an industrial scale [14,17].
2.2 PhysicalmodicationofPLGA
Modifications to the particle surface and size can be used to
avoid particle clearance, increase circulation time, improve
transport across biological barriers, or to prolong residence
time at the site of absorption. Prolonging drug presence can
be manipulated by using mucoadhesive materials, such as
the biodegradable polysaccharide chitosan. Yamamoto et al.
produced PLGA nanoparticles surface-modified with chitosan
and encapsulating the peptide elcatonin [18]. The surface-
modified nanoparticles delivered to the lungs of guinea pigs
resulted in prolonged effects compared to the unmodified nano-
particles, and the chitosan-modified nanoparticles were elimi-
nated more slowly than the unmodified version, suggesting that
nanoparticle retention was responsible for the sustained effects.
2.3 ChemicalmodicationofPLGA
Despite its many advantages, PLGA also has some inherent
shortcomings. The lack of hydrophilic and functional
groups leads to challenges regarding drug encapsulation
and stability during storage. Other challenges include
polyphasic release patterns, low encapsulation efficiency
and high burst release [19,20]. A major challenge of nano-
particle delivery to the lungs is formulation instability due
to particle–particle interactions.
A strategy to overcome some of the issues associated with
the use of PLGA for pulmonary delivery of nanocarriers was
to create a polymer that is more hydrophilic in nature than
PLGA and to introduce functional groups for improved drug–
polymer interactions within the nanoparticles. For example,
block copolymers of hydrophilic poly(ethylene glycol) (PEG)
with PLGA have been reported to show accelerated drug
release [21,22]. Star and comb-shaped PLA or PLGA could be
synthesized with multifunctional initiators, such as glycerol,
pentaerythritol, amino-propanediol, poly(vinyl alcohol) and
dextran [23-28]. A primary feature of these materials is that
they have high molecular weights but relatively short PLA or
PLGA chains, and more hydroxyl end groups, which leads to
increased hydrophilicity and faster degradation rates compared
to linear PLA or PLGA of similar molecular weight.
Polyelectrolytes with functional groups in the backbone,
such as amine and sulfonic acid groups, were introduced
into brush-like graft PLGA. These modifications affect the
colloidal stability of carrier systems by imparting positive or
negative surface charges and increasing protein or drug
loading of carriers by electrostatic interactions [29-32]. These
functional groups also accelerate the degradation rate by
enhancing the hydrophilic character of the polyester [29,31,32].
Compared with amphiphilic block copolymers, the
amphiphilic graft copolymers have multi-grafted hydrophobic/
hydrophilic branches along a hydrophilic/hydrophobic polymer
backbone. Therefore, the properties of nanoparticles can
be easily varied by simply adjusting the graft density and
side chain length of the branches.
2.3.1 PVA-PLGA
Dailey et al. reported the synthesis of a series of poly(vinyl
alcohol) (PVA)-based branched polyesters with PLGA side
chains (PVA-g-PLGA, Figure 1A) [33]. The PVA provides a
hydrophilic basis for the copolymer, while the degree of the
copolymer hydrophobicity could be varied according to
the length of the PLGA side chains grafted onto the PVA.
These copolymers exhibited a lower burst effect coupled
with a linear infusion-like release profile of proteins, which
could be controlled by the structure and molecular weight
of the copolymer. Also, in contrast to the bulk erosion
observed for PLGA, the PVA-g-PLGA copolymers exhibited
a surface erosion biodegradation mechanism [34,35]. Further
developments of this type of copolymer could satisfy the
requirements of different drugs and proteins delivered by
the pulmonary route.
PVAs have shown good protein compatibility, muco-
adhesive properties and better temperature stability during
bulk polymerization with lactide and glycolide. PVAs with
molecular weights less than 15,000 g/mol will be eliminated
from the body by renal excretion [36]. Biodegradation occurs
by surface erosion, and the biocompatibility is comparable
to that of linear PLGA [37,38].
Uncharged and charged PVA-PLGA having side chain
lengths higher than 10 have the potential for the formation
of microparticles and nanoparticles, and water soluble
polyesters with PLGA side chain lengths smaller than three
are capable of forming nanocomplexes with oppositely
charged proteins [15,27,35].
2.3.2 SB-PVA-PLGA
Varying amounts of sulfobutyl (SB) groups were attached to
the backbone to create SB-PVA-g-PLGA polymers with an
increasingly negative surface charge (Figure 1B) [39]. The
SB-PVA-g-PLGA polymer allows the preparation of
nanoparticles that exhibit a core-corona structure with the
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negatively charged hydrophilic sulfobutyl groups oriented
towards the outer aqueous phase, providing for the rearrange-
ment of the hydrophilic backbone to the particle surface.
Jung et al. demonstrated the preparation of particles with
this polymer having diameters between 100 and 500 nm [40].
These colloidal carriers can be prepared without the use
of additional surfactants. This is of extreme importance for
pulmonary application, as the inhalation of synthetic surfac-
tants may disturb the surface tension of the pulmonary lining
fluid and lead to impaired lung function or inflammation.
2.3.3 Amine-modified PVA-PLGA
Different amino groups, such as dimethylaminopropylamine
(DMAPA), diethylaminoethylamine (DEAEA) and
PVA-g-PLGA
Sulfobutyl-PVA-g-PLGA
A.
B.
C.
Amine-PVA-PLGA R′ = C3H6, C2H4; R′′ = C2H5, CH3
P(VS-VA)-g-PLGA
D.
OH S (O
*
*
O
O
O
O
OH
O
O
O
O O
O-Na+
)(
x))(
y
)(
n
z
)
m
(
N
R′′
R′′
R′
O OH (O
*
O
O
O
O
OH
O
O
O
*
O
NH
O
)((w)(
x)(
y
)(
n
)
z
)
m
*
OOH (O
*
O
O
O
O
OH
O
O
OO
)(
(x)(
y)
z
)
m
)(
n
*
OH O(O
*
O
O
O
O
OH
O
O
O
S OO
OH
)(
(x)(
y
)(
n
)
z
)
m
Figure1.ThechemicalstructureofaseriesofPVA-basedbranchedPLGAs.
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diethylaminopropylamine (DEAPA) were attached to the
backbone to create polymers with increasingly positive
surface charges (Figure 1C) [41]. Wittmar et al. developed
this class of biodegradable amine-substituted PVA
polymers, onto which hydrophobic PLGA chains were
grafted. The amphiphilic properties of diethylaminopropyl
amine-poly(vinyl alcohol)-grafted-poly(lactide-co-glycolide)
(DEAPA-PVA-g-PLGA) make it a versatile polymer for
pulmonary drug delivery [29,42]. Using a modified solvent
displacement method, nanoparticles can be generated
from DEAPA-PVA-PLGA without high shear forces.
Furthermore, degradation times can be tailored by the
degree of amine substitution to range from a few days
to several weeks, which is a vital aspect for long-term
pulmonary application. In contrast, the long degradation
rates of commercially available PLGA are a critical factor
for its use in lung delivery [43].
Coating DEAPA-PVA-g-PLGA nanoparticles with carboxy-
methylcellulose can prevent aggregation of the particles
during nebulization; Dailey et al. reported the preparation
of such particles in the size range of 76 – 256 nm [42].
Such aggregation occurs with more hydrophobic polymers.
The degradation rate decreased with increasing amounts
of CMC. DEAPA-PVA-g-PLGA nanoparticles were also
shown to be taken up into alveolar epithelial cells (A549)
to a low extent [42].
2.3.4 P(VS-VA)-PLGA
To control precisely the number of functional groups,
P(VS-VA)-g-PLGA was recently developed (Figure 1D).
The polyelectrolyte backbones were obtained by the radical
copolymerization between vinyl acetate and vinyl sulfonic
acid sodium salt and the subsequent hydrolysis. The obtained
poly(vinyl sulfonic-co-vinyl alcohol) (P[VS-VA]) backbones
were grafted with PLGA by ring-opening melt polymerization
using SnOct2 as catalyst through the hydroxyl group [44].
It was demonstrated that the degree of sulfonic acid sub-
stitution and the side length of PLGA can be easily controlled
by the feed ratio. Surface characterization studies showed
that, as observed for SB-PVA-g-PLGA, nanoparticles prepared
from these polymers exhibited a core-corona structure
with the negatively charged, hydrophilic sulfonic groups
oriented towards the outer aqueous phase [45]. Nanoparticles
prepared from this novel polymer class were reported
to range in size from 120 – 151 nm [45].
2.4 Otherpolymers
Polyanhydrides and polyacrylates have also been recently
investigated for nanocarrier-based pulmonary drug delivery
applications. Fliegel et al. describe a novel class of
biodegradable poly(ether-anhydride) polymers designed
for pulmonary drug delivery [46]. These polymers are
composed of sebacic acid and poly(ethylene glycol) (PEG)
in various ratios. By the addition of 10% PEG, the fraction
of particles deposited in the lower stages of a model
lung could be increased, most likely due to minimized
aggregation from surface roughness. Zhang et al. have
used polybutylcyanoacrylate to prepare nanoparticles for
intratracheal delivery [47].
In summary, chemical modifications to PLGA add
several benefits in the selection of a suitable material for
nanocarriers in the lung. The introduction of positive or
negative charges can enhance the encapsulation efficiency
and release profile of oppositely-charged drugs, proteins, or
genetic material. Adjustments to the polymer structure can
alter the balance between hydrophobic and hydrophilic
groups, which can in turn affect drug loading, release,
nanoparticle orientation, particle size and surface charge.
Furthermore, certain functional groups will affect the
polymer degradation rate, which not only affects the release
of active agent but is also of concern when one considers
the possible accumulation of polymer within the lung
following repeated doses. The design of fast-degrading
polymers, such as DEAPA-PVA-g-PLGA, overcomes some
of the challenges associated with polymer accumulation
due to slower degrading PLGA particles.
3. Pulmonarydeliveryofactiveagents
The previous section introduced several polymers that have
been used and designed for pulmonary delivery of nano-
carriers; this section will present some of the specific appli-
cations of these materials in the delivery of active agents,
including drugs, peptides, proteins, DNA, siRNA and vaccines.
3.1 Drugdelivery
As mentioned above, PLGA exhibits a triphasic drug and
protein release kinetic with an initial burst effect, which
is governed by diffusion kinetics, followed by a lag
phase and a secondary burst phase [48,49]. By varying
the PLGA chain lengths, the proportion of lactic to
glycolic acid and the molecular weight, drug release profiles
can be influenced [15].
Dutt et al. encapsulated isoniazid and rifampicin into
PLGA microparticles and investigated their release profile
from different formulations, demonstrating that isoniazid
shows a sustained release of up to 3 days from porous
microparticles and of up to 6 days from non-porous
microparticles. By hardening the PLGA microparticles, a
sustained release carrier system of up to 7 weeks in vitro
and in vivo could be achieved. In a murine model
one dose of PLGA microparticles was able to clear bacteria
from the lungs and liver more effectively as compared to
a daily administration of free drug [50,51].
For the treatment of tuberculosis, recombinant Myco-
bacterium tuberculosis antigen 85B (Ag85B) was encapsulated
by spray-drying into PLGA-microspheres [52]. With a median
diameter of 3 – 4 μm, these microspheres were suitable
for targeting macrophages and for aerosol delivery to the
lung. PLGA-rAg85B microspheres were able to stimulate an
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antigen response that was two orders of magnitude higher
than that observed with the pure rAG85B. In another study,
PLGA microspheres, prepared using emulsion/solvent
evaporation, were loaded with rifampicin and delivered
to guinea pigs, which were infected with Mycobacterium
tuberculosis [53]. Compared to nebulized rifampicin
suspension, the aerosolized rifampicin-loaded PLGA micro-
spheres were able to reduce most measures of tuberculosis
infection. Encapsulation of three anti-tubercular drugs –
rifampicin, isoniazid, and pyrazinamide – into PLGA
nanoparticles achieved sustained therapeutic drug levels for
6 – 8 days in the plasma, and for up to 11 days in the
lungs. The drug-loaded nanoparticles were prepared by the
multiple emulsion technique and nebulized after vacuum
drying. A significantly prolonged elimination half-life was
observed compared to the orally administered drug and
no tubercle bacilli could be detected in the lungs after
five doses of treatment [54].
A second widely used class of biodegradable polymers
for pulmonary delivery is chitosan and its derivatives.
The degradation of chitosan has already been tested in
several studies and it has been shown that chitosans are
depolymerized enzymatically by lysozyme, albeit with a very
slow rate [55-57]. Like lactoferrin or peroxide, lysozymes are
present within the lung mucus and lysozyme is the most
abundant antimicrobial polypeptide in respiratory tract
secretions [58,59]. Learoyd et al. investigated the influence of
chitosan molecular weight on the drug release of terbutaline
sulfate spray powders using low, medium and high molecular
weight chitosan. With increasing molecular weights, the
drug release profile changed from a burst release to a
sustained drug release profile over 2 – 4 h. The microparticles
generated displayed a median diameter of 1 – 2.5 μm
and were therefore suitable for inhalation [60].
Corrigan et al. investigated the influence of the preparation
media on the morphology and characteristics of chitosan
microparticles prepared by spray-drying. As the degree of
acetylation of chitosans affects its physicochemical properties
(i.e., viscosity, degradability and solubility), spray drying
was performed in hydrochloric acid or acetic acid. It was
observed that the presence of acetic acid leads to increased
acetylation of chitosan during spray-drying. When loading
chitosan microparticles with salbutamol by spray-drying,
a high respirable fraction was achieved when aerosolized
into a twin impinger. However, the burst release of the
drug in less than 5 min requires further optimization for
future pulmonary delivery [61].
3.2 Peptideandproteindelivery
With its large alveolar surface area, thin epithelial barrier
and low proteolytic activity compared to other administration
routes, the lung represents an attractive route for the delivery
of macromolecules, such as proteins. Due to their extreme
sensitivity, the design of sophisticated drug carriers is
required to overcome the many barriers of the lungs.
Amidi et al. generated insulin-loaded microparticles by
spray-drying using N-trimethyl chitosan [62]. In all formu-
lations the secondary and tertiary structure of insulin could
be preserved. Even after 1-year storage at 4°C, the particle
characteristics and insulin structure remained unchanged
and intact. Grenha et al. developed a microparticulate
carrier system for insulin-loaded chitosan nanoparticles.
Using mannitol and lactose as excipients, the insulin-
loaded chitosan nanoparticles were microencapsulated
by spray-drying, yielding particles with Ferret diameters of
2 – 4 μm. In vitro studies showed that approximately
75 – 80% of the encapsulated insulin could be released from
the nanoparticle-loaded microspheres within 15 min [63].
PLGA nanospheres coated with chitosan for pulmonary
delivery of the peptide elcatonin have been mentioned
previously as an example of the advantages of physical
modifications to PLGA nanocarriers [18].
Kawashima et al. dosed PLGA nanoparticles prepared
with insulin to guinea pig lungs and demonstrated a
significant reduction in blood glucose level, with a prolonged
effect over 48 h compared to insulin solution [64].
Insulin-loaded nanoparticles using a different polymer,
poly(butyl cyanoacrylate), delivered to the lungs of rats,
were shown by Zhang et al. to extend the duration of
a hypoglycemic effect over 20 h [47].
3.3 Gene,siRNAandvaccinedelivery
An ideal gene delivery system should show high transfection
levels, be non-toxic and biodegradable for long-term appli-
cation. Polyethylenimine (PEI) is one of the most effective
cationic compounds for plasmid delivery into mammalian
cells [65,66]. The cationic groups of carriers such as PEI
can form complexes with oppositely-charged genetic cargo.
However, the high toxicity and lack of biodegradability
of PEI limits its potential for pulmonary application.
To overcome these drawbacks of PEI and at the same
time maintain its high transfection efficiency, Thomas et al.
developed biodegradable PEIs composed of a linear 423 Da
PEI and a branched 1.8 kDa PEI [67]. These two low
molecular weight PEIs, which have been shown to be
less toxic than its high molecular weight counterpart [68],
are crosslinked with bi- and oligo-functional acrylates
to obtain biodegradable high molecular weight PEIs.
Creating a combinatorial library of vectors, it was
shown that the optimal vector in vivo was the mixed PEI
crosslinked with propylene glycol glycerolate diacrylate. It
combined the highest transfection efficiency (186 times
higher than the physical mixture of the parental PEIs)
with low toxicity, whereas the commercially available 22-kDa
PEI caused 50% mortality.
Another group of biodegradable polymers used for gene
delivery are chitosans, which are considered to be non-toxic.
One of their major drawbacks, however, is the modest
transfection efficiency in vivo and in vitro [69,70]. Further
investigations are necessary to clarify the mechanisms of
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uptake and transfection efficiency. Köping-Höggard et al.
administered chitosan-pDNA polyplexes to the lungs of mice,
showing that the high molecular weight chitosan is well
tolerable to mice and is able to promote gene delivery into
the lungs. However, PEI remained a superior pDNA vector
for pulmonary gene delivery [71]. After further optimization,
a comparable luciferase gene expression to that of PEI was
achieved when administered to mouse lungs. The molecular
weight of chitosans in this study ranged from 1.2 – 10 kDa.
Physicochemical studies showed that low molecular weight
chitosans are able to release pDNA in the presence of the
model anion heparin, leading to a higher transfection
efficiency compared to the more stable high molecular
weight chitosan polyplexes [72].
Another approach used to improve the bioactivity of
chitosan oligomer polyplexes was to introduce a trisaccharide
branch that targets cell surface lectins [73]. Lectins
are reported to be exposed on airway epithelial cells
and have the ability to bind sugar residues [74,75]. The trans-
fection efficiency of the trisaccharide-substituted chitosans
was significantly higher in human liver hepatocytes (HepG2)
and a human bronchial epithelial cell line (16HBE14o-)
than unmodified chitosan and PEI 25 kDa. Luciferase
gene expression in mouse lungs was fourfold higher for
trisaccharide-substituted chitosans than unmodified chitosan
but unfortunately no comparison to PEI 25 kDa in vivo
was shown.
Chitosan has also been used for siRNA delivery to
human lung carcinoma cells (H1299). With a chitosan/
siRNA formulation containing sucrose as lyoprotectant, a
70% knock down was achieved [76]. Howard et al. prepared
polyplexes between siRNA and chitosan [77]. Effective
knockdown, both in vitro and in vivo, was observed. In
the H1299 human lung carcinoma cell line and in murine
peritoneal macrophages, knockdowns of 77.9 and 89.3%,
respectively, were achieved. With a 40% reduction in
EGFP fluorescence in bronchial epithelial cells of transgenic
EGFP mice, chitosan/siRNA polyplexes showed successful
RNA interference.
Chitosan has also been reported to display therapeutic
potential in the case of respiratory syncytial virus (RSV).
RSV causes bronchiolitis and pneumonia and is also a severe
risk factor for asthma. Treatment of rats prior to RSV
infection with chitosan/siRNA-polyplexes containing a
siRNA that interferes against the RSV-NS1 gene (siNSI),
reduced the virus titers in the lung. SiNS1-treated rats
showed less inflammation and hyperresponsiveness compared
to the control [78].
Pulmonary DNA vaccination represents a non-invasive
and less painful administration route for immunization. The
opportunity to combine the genetic information of various
antigen epitopes and cytokines, easy production and the
high stability of plasmid DNA compared to recombinant
proteins and pathogens make it an attractive class of vaccines.
Many pulmonary pathogens, such as M. tuberculosis bacillus,
respiratory syncytial virus (RSV) and severe acute respiratory
syndrome corona virus (SARS) could all be treated
once a suitable vaccination has been developed [79-81].
Bivas-Benita et al. prepared poly(D,L-lactide-co-glycolide)
(PLGA) nanoparticles and coated them with polyethylenimine
(PEI). These PLGA-PEI nanoparticles were loaded with
DNA and uptake into the endo-lysosomal compartment
of the human airway submucosal epithelial cell line,
Calu-3, was detected [82].
N-trimethyl chitosan and dextran microparticles were
investigated for pulmonary delivery of diphtheria toxoid.
The microparticles were prepared by drying an aqueous solution
of polymer and diphtheria toxoid using a supercritical fluid
(SCF) spraying process. In contrast to dextran microparticles,
only the N-trimethyl chitosan microparticles led to a detectable
secretion of IgA when administered to the lungs [82].
4. Conclusions
Pulmonary drug delivery is attractive for both local and
systemic drug delivery as a non-invasive route that provides
a large surface area, a thin epithelial barrier, high blood flow
and the avoidance of first-pass metabolism. Nanoparticles
have several advantages for controlled drug delivery by the
pulmonary route, including sustained release, reduced dosing
frequency, as well as being an appropriate size for avoiding
alveolar macrophage clearance or promoting transepithelial
transport. Particles or nanoparticle clusters with aerodynamic
diameters between 1 and 5 μm have the highest probability
of successful lung deposition. The selection of natural or
synthetic biodegradable polymeric materials for nanocarriers
is important in order to design particles with the desired
characteristics. Biocompatibility, size, charge and drug release
rates must all be considered, but in order to avoid accumu-
lation of polymeric materials following repeated dosing, the
polymer degradation rate is crucial.
5. Expertopinion
Although significant progress has been made in recent years
relating to the design of biodegradable nanocarrier strategies
for the delivery of drugs, peptides, proteins, genes, siRNA
and vaccines, the future of pulmonary delivery strategies is
expected to be influenced critically by the outcome of
ongoing discussions.
An issue that remains surrounded by considerable debate
is the question whether the lung should be used as an entry
port for systemic drug administration. In this context, the
safety of the nanocarriers and a lack of inflammatory and
immunogenic potential need to be demonstrated under
chronic treatment conditions. Such studies have not been
presented with drug-loaded nanocarriers, but will be
necessary during future clinical trials. Such issues are not
limited to pulmonary drug delivery, but are also important
in oral and intravenous administration.
Biodegradablepolymericnanocarriersforpulmonarydrugdelivery
8 ExpertOpin.DrugDeliv.(2008) 5(6)
647
650
655
660
665
670
674
619
625
630
635
640
646
A second area where a general lack of information can be
recognized is the interface between nano-objects and lung
tissue/cells. There are still many questions to be answered
with regard to the fate of nanocarriers in the lung. For
example, what physical and chemical characteristics affect
clearance by alveolar macrophages? Which particle properties
affect cellular internalization and transport across the
pulmonary epithelium? Furthermore, predictive correlations
between in vitro, ex vivo and in vivo models are necessary in
preparation for clinical trials of nanocarrier-based drug
delivery systems in the lung.
The third area where more fundamental information
needs to be generated is related to the aforementioned topics
but addresses the aspects of biomaterials used for pulmonary
delivery systems. The residence time of nanocarriers in lung
tissue, their degradation mechanisms and the clearance of
degradation products will ultimately affect the safety and
biocompatibility of such delivery devices. The accumulation
of carrier materials within the lung, including polymer and
its degradation products, may bring about long-term concerns
that outweigh the benefits of therapy with polymeric carriers.
More polymers with short half-lifes would clearly be
desirable, as this would be more suitable for repeated admini-
stration. A consensus about testing strategies, including both
in vitro and ex vivo models, has yet to be reached.
The fourth area where advances would be desirable relates
to the design of nanocarriers, especially the incorporation of
sensitive therapeutic agents, such as proteins, p-DNA and
siRNA, which require particle production methods avoiding
high shear stress. Important progress regarding particle
stabilization is necessary to prolong the shelf life of
nanoparticles intended for pulmonary delivery. Reconstitution
of dried nanosuspensions for nebulization or packaging of
nanoparticles into adequately-sized particle clusters for
inhalation of a dry powder are important steps to deliver the
desired doses to the desired regional targets within the lung.
As diffusion distances in nanocarriers are shorter by
definition, control over drug loading and release under
in vitro as well as in vivo conditions remains a challenge. In
addition to new technologies, polymer design may also help
to address such problems, as we have shown in some of the
examples presented above.
Pulmonary drug delivery is a fascinating area of research
which needs input from various disciplines ranging from
medical sciences to aerosol physics. As these interdisciplinary
research activities continue in the area of biodegradable
nanocarriers for pulmonary drug delivery, one can expect
significant advancements in the future that will extend hope
to healthcare professionals and patients alike.
Acknowledgements
The authors wish to thank the German Ministry for
Education and Research (BMBF) for a nanotechnology
science award 13N8889 and Boehringer Ingelheim
Pharma GmbH for technical support.
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Afliation
Erik Rytting1 PhD, Juliane Nguyen2,
Xiaoying Wang3 PhD & Thomas Kissel†4 PhD
†Author for correspondence
1Postdoctoral Research Associate
Philipps-Universität Marburg,
Institut für Pharmazeutische
Technologie & Biopharmazie,
Ketzerbach 63, D-35032 Marburg, Germany
2PhD candidate
Philipps-Universität Marburg,
Institut für Pharmazeutische
Technologie & Biopharmazie,
Ketzerbach 63, D-35032 Marburg, Germany
3Postdoctoral Research Associate
Philipps-Universität Marburg,
Institut für Pharmazeutische
Technologie & Biopharmazie,
Ketzerbach 63, D-35032 Marburg, Germany
4Professor and Head
Philipps-Universität Marburg,
Institut für Pharmazeutische
Technologie & Biopharmazie,
Ketzerbach 63, D-35032 Marburg, Germany
Tel: +49 6421 2825881;
Fax: +49 6421 2827016;
E-mail: kissel@staff.uni-marburg.de
