Merging the best of both worlds: hybrid lipid-enveloped matrix nanocomposites in drug delivery
Koen Raemdonck*,a, Kevin Braeckmansa,b, Jo Demeestera and Stefaan C. De Smedt**,a
a
Ghent Research Group on Nanomedicines, Laboratory of General Biochemistry and Physical
Pharmacy, Faculty of Pharmaceutical Sciences, Ghent University, Harelbekestraat 72, B-9000 Ghent,
Belgium
b
Center for Nano-and Biophotonics, Ghent University
*Corresponding author during submission: Dr. Koen Raemdonck
koen.raemdonck@ugent.be
Tel: +32 9 2648078
Fax: +32 9 2648189
**Corresponding author after submission: Prof. Dr. Stefaan C. De Smedt
stefaan.desmedt@ugent.be
Laboratory of General Biochemistry and Physical Pharmacy
Faculty of Pharmaceutical Sciences
Ghent University
Harelbekestraat 72, 9000 Ghent, Belgium
Tel: +32 9 2648076
Fax: +32 9 2648189
1
ABSTRACT
The advent of nanotechnology has revolutionized drug delivery in terms of improving drug efficacy
and safety. Both polymer-based and lipid-based drug-loaded nanocarriers have demonstrated clinical
benefit to date. However, to address the multifaceted drug delivery challenges ahead and further
expand the spectrum of therapeutic applications, hybrid lipid-polymer nanocomposites have been
designed to merge the beneficial features of both polymeric drug delivery systems and liposomes in a
single nanocarrier. This review focuses on different classes of nanohybrids characterized by a drugloaded polymeric matrix core enclosed in a lipid shell. Various nanoengineering approaches to obtain
lipid-polymer nanocomposites with a core-shell nanoarchitecture will be discussed as well as their
predominant applications in drug delivery.
1. INTRODUCTION
Nanotechnology has found widespread use in pharmaceutical and biomedical applications, including
improved delivery of small molecule drugs, macromolecular therapeutics and/or imaging agents. To
improve the overall therapeutic index of existing low molecular weight therapeutics, their encapsulation
in nanosized carriers has demonstrated great benefit. Likewise, to attain the full therapeutic potential of
novel biopharmaceuticals such as proteins, peptides and nucleic acids, formulating them in
nanoparticles (NPs) is generally required.1-10 Rational design of drug-loaded nanocarriers, so-called
nanomedicines, may aid in overcoming physicochemical limitations associated with free drugs, i.e.
solubility and stability in aqueous media. NPs should additionally allow facile encapsulation of drug
cargo with high efficiency and prevent premature drug release. Ultimately, NPs ideally encompass
multiple functionalities devised to overcome the various extra-and intracellular barriers imposed by
the human body. The latter implies avoiding fast renal clearance, protecting the active payload
against
(enzymatic)
degradation,
bypassing
the
body’s
immune
defenses,
improving
pharmacokinetics toward the diseased tissue and, in case an intracellular target is envisioned,
actively promoting intracellular drug delivery.1-3,8,11 If successful, nanomedicines may thus enhance
2
the fraction of the administered dose that effectively reaches the intended target site and mitigate
potential off-target toxicity. Moreover, in the light of personalized medicine, flexible design of
targeted drug carriers that enable tailored controlled release of drugs or synergistic drug
combinations with distinct physicochemical properties could provide a definite asset.12
The dominant types of conventional nanomedicines to date, i.e. liposomes 7,13 and polymeric NPs 2,8,
generally have failed to combine all of the complex requirements outlined above. Liposomes can be
defined as self-assembled vesicles comprising one or multiple concentric lipid bilayers that enclose
an aqueous core. The versatility of liposomes is reflected in their ability to carry both hydrophilic and
hydrophobic drugs in the aqueous lumen and lipid bilayer respectively, their outstanding safety
profile and ease of surface functionalization with hydrophilic polymers like poly(ethylene glycol)
(PEG) and/or ligands to obtain long-circulating targeted nanomedicines. Decades of research has
converged in the clinical application of liposomal formulations for a wide variety of drug
molecules.7,13 In spite of the many favorable characteristics they exhibit, liposomes may experience
low loading efficiencies, especially for hydrophobic drugs. In addition, dependent on their
composition, liposomes may suffer from poor stability in vivo, resulting in unwanted burst drug
release.13-15 Polymeric NPs may provide a valuable alternative to overcome some of the abovementioned limitations, as they generally demonstrate outstanding drug loading capacity for drugs
and/or contrast agents with diverging physicochemical properties. Moreover, progress in polymer
chemistry enables the design of NPs with ample control over their nanoarchitecture and biophysical
properties which facilitates application in controlled and triggered drug release strategies. On the
other hand, issues have been raised on the biocompatibility of certain polymeric materials and the
potential heterogeneity and low density of chemical surface functionalization. 2,8,16-18
To address the multifaceted drug delivery challenges outlined above, several research groups have
turned to the design of hybrid lipid-polymer nanocomposites with the primary aim to combine the
most valuable features of both polymeric drug delivery systems and liposomes in the design of lipidpolymer nanocomposites.14,19-22
3
Many classes of lipid-coated nanomaterials have been described in the literature. For example,
lipopolyplexes can be defined as self-assembled ternary electrostatic complexes comprising a
nanosized polyanion:polycation complex (polyplex) subsequently coated with (oppositely charged)
lipids.23,24 The lipid bilayer/multilayer coat confers additional colloidal stability to the ensemble and
serves to incorporate key functional moieties, as demonstrated in numerous publications by the
groups of Leaf Huang
25-29
and Hideyoshi Harashima
30-34
. This type of hybrid nanocomposites has
most often been employed to encapsulate and deliver nucleic acid therapeutics. A remarkable
example, highlighting the drug delivery potential of these hybrid NPs for siRNA, has been put forward
by Peer et al.. They succeeded in designing multifunctional antibody-targeted lipopolyplexes to direct
anti-inflammatory siRNA toward gut-associated leucocytes following systemic administration to
tackle experimentally induced colitis in mice.23 Furthermore, in the emerging field of
nanotheranostics, aiming to merge both therapeutic and imaging modalities in a single
nanostructure, inorganic NPs like semiconductor nanocrystals (quantum dots) and magnetic NPs
have been provided with a lipid coat to optimize their in vivo application in fluorescence or magnetic
resonance imaging (MRI) respectively.35-37 Although of outstanding interest for drug delivery, the
nanocomposites described above are beyond the scope of this review. Here, we will predominantly
focus our attention on lipid-enveloped polymer matrix nanoparticles, typically defined by a core-shell
architecture in which a (drug-loaded) matrix core is surrounded by a lipid shell. Polymeric matrix
nanoparticles can be defined as colloidal 3D polymer networks with a size ranging from 10 nm to 1
µm, in which the drug can be physically complexed/dispersed/dissolved or chemically coupled to the
polymer chains.3 In contrast to conventional polyplexes, in a matrix NP the encapsulated therapeutic
compound typically does not structurally contribute to NP formation.38
The consequential advantages of combining lipids and polymers in a hybrid drug delivery platform
are situated on many levels.14,19-22 Efficient drug encapsulation can be achieved both in the polymeric
core as in the surrounding lipid envelope. Drug release can be controlled by polymer degradation but
also modulated by the presence of the lipid coat that acts as a diffusional barrier. In this way,
4
unwanted drug leakage from the nanocarrier can be prevented. Furthermore, the lipid layer may
hinder the influx of water, tempering polymer hydrolysis and slowing down drug release. The lipid
shell may impart many other valuable traits such as the ease of incorporating PEGylated lipids and
targeting ligands, the obscuration of toxicity associated with the polymer core, the improvement of
colloidal stability and enhancement of intracellular drug delivery. In addition, the polymer matrix
core may also contribute to the structural integrity of the lipid coat.
Our major aim is to discuss various inspiring reports on lipid-polymer nanocomposites that
successfully implemented this synergistic drug delivery approach, focusing on lipid-coated solid
hydrophobic polyester NPs, mesoporous silica NPs (MSNPs) and hydrophilic hydrogel nanoparticles
(nanogels). Different preparation methods to obtain lipid-polymer nanocomposites with a core-shell
nanoarchitecture will be briefly described and novel nanoengineering approaches will be highlighted.
Finally, we present an overview of drug delivery applications, mainly situated in the context of cancer
therapy and vaccination.
2. GENERAL CONSIDERATIONS ON NANOMEDICINAL DRUG DELIVERY
Upon instillation of NPs in a biological medium or in contact with target cells, the size, shape and
surface properties of the NP will largely define its factual behavior at the so-called nano-bio
interface.39-41 With regard to the polymer-lipid core-shell NPs described here, this implies that the
external lipid shell (and the modifications it harbors) will significantly determine the dynamic
interactions with the myriad of biological components the hybrid NP will encounter. Following
parenteral administration, in general unmodified NPs are rapidly removed from the blood circulation
via the mononuclear phagocytic system (MPS), by which NPs will predominantly accumulate in liver
and spleen.40-42 The latter is facilitated in a passive manner via NP extravasation through the
sinusoidal capillaries in these organs that are characterized by a fenestrated discontinuous
endothelium and in an active fashion by the deposition of particular blood proteins that opsonize the
NP surface and mediate recognition and clearance by hepatic/splenic phagocytes. Interaction with
5
circulating proteins can additionally evoke NP aggregation, causing them to be trapped in the lung
capillary bed.40 To reduce MPS clearance and maintain colloidal stability, it has become common
practice to shield the NP surface with a hydrophilic stealth layer, most commonly composed of
poly(ethylene glycol) (PEG), which is registered by the US Food and Drug Administration (FDA) as a
GRAS polymer (“Generally Recognized As Safe”).2,43 In this way, the blood circulation time can be
significantly extended, which is of particular interest to augment the passive extravasation of
nanomedicines in solid tumors and sites of inflammation, as these are characterized by an increased
vascular permeability. Besides the defective angiogenesis in cancerous tissues, accounting for the
leaky vasculature, also the lymphatic drainage is (partially) impaired thus contributing to NP
accumulation in the tumor interstitium. This so-called enhanced permeation and retention (EPR)
effect may allow the infiltration of nanosized drug carriers up to ~400 nm in diameter,1,44 although
this is highly dependent on tumor type and microenvironmental factors (Figure 1). Even within the
same tumor, the endothelial permeability may be quite heterogeneous.2,40,45 In general, to
experience optimal EPR, NPs should be larger than 10 nm to avoid glomerular filtration by the
kidneys and should not exceed 200 nm in size.1,41 Nevertheless, it is known that within this size range
the NPs are also able to penetrate into the liver tissue.10
In the situation where an extracellular molecular target is envisioned, controlled release of the active
compound in the target tissue interstitium may suffice for therapeutic benefit. However, many drugs
require delivery across the cellular barriers. Hydrophobic small molecule anticancer drugs, when
released from their nanocarrier in the tissue extracellular matrix, are able to passively diffuse across
the plasma membrane of diseased cells to interact with their intracellular target. On the other hand,
this uncontrolled arbitrary nature of drug delivery may hamper maintaining active tissue drug levels
within an acceptable range and could promote the development of a multi-drug resistant (MDR)
phenotype.1,46,47
Cancer cell drug resistance is dominated by the activity of MDR membrane
transporter proteins that mediate efflux of a variety of chemotherapeutic drugs. Cancer cells that
lack drug efflux mechanisms or only moderately express these transporters will preferentially be
6
affected by the drug, eventually promoting selective survival of a drug resistant cell population.
Altogether, these claims advocate targeted intracellular drug delivery via nanomedicines to increase
intracellular drug concentrations and bypass efflux pump mediated drug resistance 1,10 (Figure 2).
Figure 1. Combined passive and active tumor targeting. Enhanced tumor tissue accumulation of
nanomedicines is achieved via the leaky vasculature and impaired lymphatic drainage, i.e. the
enhanced permeation and retention effect (EPR). Decorating the nanomedicines with targeting
ligands allows the specific binding of tumor cells followed by receptor-mediated internalization. Drug
release can occur in the extracellular environment or directly in the cell after endocytic uptake. The
latter delivery method leads to higher intracellular drug concentrations that can bypass efflux-based
drug resistance. (Reprinted by permission from Macmillan Publishers Ltd: Nature Nanotechnology1,
Copyright 2007).
7
Figure 2. Overview of the predominant intracellular barriers encountered in nanomedicinal drug
delivery. Following endocytic uptake, nanomedicines are confined within endosomal vesicles. Escape
from the endosomes is required for drugs that need to be delivered in the cell cytoplasm or
nucleoplasm and to avoid endolysosomal degradation (e.g. for nucleic acid therapeutics). Small
lipophilic drugs can passively permeate across the plasma membrane or endosomal limiting
membrane when released in the extracellular matrix (ECM) or endosomal lumen, respectively.
Cancer cells can adopt a multidrug-resistant (MDR) phenotype by expressing membrane transporters
(e.g. p-glycoprotein) that actively expel various chemotherapeutics, thus lowering the intracellular
drug concentration.
Likewise, the novel class of biopharmaceuticals mainly comprises large, hydrophilic and charged
molecules (e.g. peptides and nucleic acids) that require the aid of an optimally designed nanocarrier
to reach the intracellular site-of-action. It is well-established that nanomedicines as a rule are
internalized via distinct endocytic entry portals as a function of cell type and NP physicochemical
properties.5,6,48,49 For instance, NP size is an important regulator that influences cellular uptake and
subsequent intracellular trafficking.48 Larger NPs (>500 nm) are preferentially engulfed via
phagocytosis by specialized cell types of the immune system, i.e. macrophages, monocytes and
dendritic cells.48,50 Of note, in (cancer) vaccination or anti-inflammatory treatments these cells may in
fact be the desired targets for drug delivery when aiming to modulate local immune responses and
8
incite humoral and/or cellular (anti-tumor) immunity.51-55 Smaller NPs, in the acceptable size range
for in vivo biodistribution, can be internalized via a clathrin-mediated or clathrin-independent
mechanism.48,50 On the one hand, cell uptake can be triggered by non-specific hydrophobic or
electrostatic interactions with the target cell membrane. Alternatively, surface functionalization of
NPs with bioligands, such as antibodies 53,56, peptides 57,58 or aptamers 59-61, may allow for cell-specific
recognition and internalization via receptor-mediated endocytosis. Following endocytic uptake, drugloaded NPs tend to accumulate in endolysosomes, packed with acid hydrolases, or are prone to
exocytosis in which the luminal content of multivesicular late endosomes is again expelled into the
surrounding medium.62 To avoid the endolysosomal degradation pathway, many strategies to
promote drug delivery into the cytosol have been reported in the literature. Small lipophilic drugs,
able to passively permeate across the endosomal limiting membrane, can escape from the
endosomal compartment unaided.2 For cytosolic delivery of (large) hydrophilic compounds,
nanocarriers have to rely on their intrinsic pH buffering capacity, osmotic swelling and/or fusogenic
activity to disrupt endosomal membrane integrity.6,63 A very recent report, aiming to unravel the
intracellular processes that govern the efficiency of liposome-mediated small interfering RNA (siRNA)
delivery, demonstrated that prolonging the intracellular retention of the NPs in late endosomes
substantially improved RNA interference (RNAi) mediated gene silencing.62 This study suggests that
strategies able to bypass standard endocytic recycling, might prove beneficial toward cytosolic drug
delivery, especially when endosomal escape is the main limiting factor. Altogether, it is of key
importance to combine a detailed understanding of the underlying extra-and intracellular barriers
together with innovations on the interface of materials chemistry, nanotechnology, pharmacy and
medicine.48,64-66 Novel insight in the biological behavior of nanomedicines related to successful drug
delivery should fuel the rational design of next generation nanomedicines in order to unleash the full
benefits of nanomedicinal drug delivery. The favorable features associated with lipid-polymer matrix
nanocomposites for drug delivery across the many barriers and advances in lipid-polymer
9
nanoengineering might therefore act in concert to accelerate the lab-to-clinic transition of
nanomedicinal products.
3. LIPID-POLYMER NANOENGINEERING
Given the often complex composition of lipid-polymer nanocomposites and the broad range of
techniques that is available for their production, a myriad of hybrid nanoformulations can be
constructed. The properties of the bulk material and the applied synthesis approach will largely
govern the physicochemical properties of the resulting NPs (e.g. size, stability, drug encapsulation
and release) and hence also their biological performance as a drug delivery carrier. Broadly, the
currently available synthesis techniques can be classified in two main categories, i.e. a two-step and
single-step approach.19 Both can be defined as bottom-up approaches, taking advantage of the
inherent properties of the building blocks to obtain the desired supramolecular structure.67
Commonly, a two-step synthesis involves the formation of polymeric matrix NPs and liposomes via
separate protocols after which they are merged into lipid-polymer core-shell nanohybrids in a final
step. Importantly, a two-step synthesis also implies that the size and shape of the resulting
nanohybrid will be largely governed by that of the polymeric template. Although multicomponent
hybrid NPs may definitely expand the performance of traditional nanomedicines, one of the major
limitations is inadequate control over the production process. For this reason, ample research effort
has also gone toward the optimization of single-step protocols, mostly based on a controlled selfassembly of the various molecular components in a one-pot synthesis. Both preparation methods, if
available, will be detailed below for lipid-coated polyester NPs, MSNPs and nanogels, prior to
discussing potential drug delivery applications in section 4.
3.1. Lipid-coated polyester nanoparticles
Synthetic polymeric NPs, e.g. constructed from the polyesters poly(lactic acid) (PLA) and its
copolymer with poly(glycolic acid), i.e. poly(lactic-co-glycolic acid) (PLGA), have been studied
10
extensively for drug delivery purposes the last few decades.2,3,8,68-70 Their biodegradable and
biocompatible features largely account for their widespread use in biomedical applications and
justify the approval of PLGA micro-and nanoparticles for human use in drug delivery by the FDA and
the European Medicine Agency (EMA).2,69 PLGA and PLA NPs slowly degrade via surface erosion in an
aqueous environment due to hydrolysis of the ester linkages in the (co)polymer backbone, to
produce the original monomers lactic acid and glycolic acid. The latter can be further metabolized in
the human body, which explains the favorable toxicity profile of these polyesters.71 A multitude of
therapeutic molecules has already been encapsulated in the polymer matrix, ranging from small
hydrophobic chemotherapeutics to large hydrophilic macromolecules, such as proteins and nucleic
acids. The kinetics of drug release from PLGA/PLA-based matrices is mainly governed by polymer
matrix degradation and drug diffusion and can be controlled over periods ranging from hours to
weeks.70 Conjugating drug molecules directly to the PLA/PLGA chains allows to prolong drug release
as a function of ester hydrolysis without the risk of premature diffusive drug release. The latter might
be of particular value for sub-100 nm particles that are endowed with a large surface-to-volume ratio
and are typically associated with a high risk of substantial burst release of drug molecules adsorbed
on the NP surface.72,73
11
Figure 3. The most common techniques for the preparation of polymeric nanoparticles are (A)
emulsion-solvent evaporation and (B) nanoprecipitation. The former method involves the
emulsification of a non-water miscible organic phase (containing polymer and hydrophobic drugs) in
an aqueous phase supplemented with appropriate surfactants. Subsequent removal of the organic
solvent transforms the nanoscopic emulsion droplets in solid drug-loaded polymeric NPs. In contrast,
the nanoprecipitation technique requires the dissolution of polymer and drug in a water-miscible
organic solvent that is added dropwise into the water phase (the non-solvent). Nanoprecipitation
occurs through rapid solvent diffusion and subsequent evaporation. (Reproduced with permission
from ref 3. Copyright 2012 The Royal Society of Chemistry).
A variety of methodologies is available to fabricate polymeric NPs, commonly based on the selfassembly of preformed (co)polymers via an emulsion-solvent evaporation or nanoprecipitation
technique (Figure 3), in which the latter usually yields the smallest sized NPs.2,3,69,70 Both techniques
typically lead to spherical matrix NPs. Of note, DeSimone’s group reported on an alternative soft
lithographic technique termed PRINT (Particle Replication in Non-wetting Templates) to engineer
monodisperse submicron PLGA particles of distinct yet well-defined sizes and shapes (Figure 4).74,75
The PRINT process employs perfluorinated polyether elastomeric molds for imprint lithography in
which the shape of each individual particle is determined by the cavities present in the mold.76
12
Figure 4. Particle Replication in Non-wetting Templates (PRINT) allows fabrication of monodisperse
PLGA nano-and microparticles of well-defined size and shape. Scale bars (A) 5 µm, (B) 4 µm, (C) 3 µm,
(D) 10 µm, (E) 3 µm, and (F) 20 µm. (Reprinted with permission from ref 74. Copyright 2011 American
Chemical Society).
Multiple functionalities can be incorporated into polymeric NP design
77
to assist in drug
encapsulation, drug release, in vivo biodistribution, cellular targeting and intracellular trafficking,
which has recently been comprehensively reviewed by others.2,3,8 One particular strategy to
modulate the surface of polymeric NPs, and hence also influence their behavior at the nano-bio
interface, is by depositing a stabilized (phospho)lipid layer onto the hydrophobic NP core. As for
liposomal NPs, this lipid layer can be employed to anchor amphiphilic PEGylated lipids to the surface
of the particles and/or conjugate targeting moieties. However, another important motivation to
invest in the development of these more complex lipid-polymer hybrids lies within their anticipated
improved in vivo stability, increased drug loading and added control over the drug release
process.2,19-21 The most straightforward approach to obtain lipid-polymer NPs with a core-shell
morphology is by mixing preformed polymeric NPs with preformed liposomes at the desired vesicle
to nanoparticle ratio.78-81 Electrostatic interaction between the anionic PL(G)A surface and cationic
13
liposomes is frequently applied as driving force behind the lipid coating process.79,80 Unilamellar
liposomal vesicles can easily be prepared via extrusion, sonication or high pressure homogenization.7
The same techniques are often also applied to the NP-liposome mixture to provide additional energy
input to the mixing process and facilitate lipid reorganization and fusion onto the NP surface.
Alternatively, drug loaded core NPs can also be used to directly hydrate a preformed lipid film,
followed by an appropriate sizing method, to obtain lipid-coated polymer hybrids.73,82,83 Irrespective
of the applied method, the mixing of polymeric NPs and lipids usually occurs at a temperature
exceeding the gel-to-liquid phase transition temperature of the lipid.
The production of lipid-polymer hybrid NPs via two (or more) consecutive steps may lead to marked
batch-to-batch variation, which could conceivably alter the drug delivery performance and thus
hamper clinical translation. Many research groups therefore invested in the optimization of singlestep formulation strategies, allowing a more feasible and reproducible synthesis of lipid-polymer
hybrids with better control over the physicochemical characteristics of the final hybrid NP construct.
The single-step approach implies the mixing of polymer and lipid solutions prior to nanoprecipitation
or emulsion-solvent evaporation to self-assemble the lipid-coated polymer NPs. Herein the used
phospholipids themselves act as stabilizer instead of the conventionally used poly(vinyl alcohol)
(PVA) or polypropylene oxide-polyethylene oxide block copolymers.2,84-86 Depending on the used
preparation method, solid NPs with lipid monolayer, bilayer or multilayer are produced. The group of
Farokhzad optimized the single-step self-assembly of drug-loaded PLGA NPs encapsulated within a
lecithin lipid monolayer containing interspersed PEGylated phospholipids via a modified
nanoprecipitation method. To this end, an acetonitrile solution of PLGA and hydrophobic drug was
mixed with an aqueous ethanolic (PEGylated) lipid dispersion, preheated above the lipid transition
temperature. The mixture was subsequently stirred for 2h to allow self-assembly of lipid-coated
PLGA NPs via solvent-diffusion and evaporation of the organic solvent.60,87 The presence of a lipid
monolayer resulted in reduced drug release rates as a function of lipid coverage, as demonstrated for
14
the small hydrophobic anticancer drug docetaxel, likely due to a delayed drug diffusion across the
lipid layer and a reduced water penetration inside the PLGA core slowing down ester hydrolysis.60,87
PEGylated lipid-polymer NPs can also be fabricated using PEGylated amphiphilic polyester as
precursor material instead of applying PEGylated lipids. Cationic lipid-polymer hybrid NPs were thus
produced via a single-step nanoprecipitation of cationic lipid with methoxy-poly(ethylene glycol)block-poly(lactide) (mPEG-PLA). The resulting particles consist of a hydrophobic PLA core with a nonfouling PEG shell and a monolayer of cationic lipid at the core-shell interface.88 In contrast with the
lipid-monolayer described above obtained via nanoprecipitation, Bershteyn et al. could clearly
visualize lipid-bilayer coated particles via cryo-TEM imaging, obtained with their optimized emulsionsolvent evaporation protocol. The authors could show that the finally obtained surface structure of
the hybrid NPs - distinguishing single bilayer shell, multilayer onion or flower-like lipid configurations
- was highly dependent on the amount of lipids used and the lipid composition.89 However, a major
impediment of the nanoprecipitation and emulsion-solvent evaporation technique is the poor
encapsulation of hydrophilic therapeutics, e.g. pharmaceutical proteins and peptides, which can be
resolved by virtue of a double (water-in-oil-in-water) emulsification process.2,3,69 Encapsulation of
ovalbumine (OVA) as a model protein antigen in lipid-coated PLGA NPs obtained via a doubleemulsion solvent evaporation self-assembly process was successfully demonstrated by Stephan and
coworkers. In comparison with liposomes of equal composition, the lipid-coated PLGA NPs could
encapsulate higher amounts of the OVA peptide and showed markedly delayed release kinetics. 90
From a manufacturing perspective, the combination of multiple molecular components with distinct
physicochemical properties into complex nanocomposites could additionally obstruct transition from
lab-scale to large-scale synthesis. To facilitate the scaling up of the fabrication process, various
nanoengineering approaches have recently been explored. For instance, a single step sonication
method was proposed in order to render the nanoprecipitation production process less laborious
and time-consuming. By replacing the commonly applied heating, vortexing and solvent evaporation
15
steps
60
with a single bath sonication step, the required production time to obtain lipid-monolayer
coated PLGA NPs via nanoprecipitation could be significantly reduced without affecting their
physicochemical properties.91 Nevertheless, this approach is only practicable at lab-scale and batchto-batch variations in particle synthesis still cannot be excluded. In response to these constraints, the
same authors described the production of lipid-polymer hybrids via a continuous flow-confined
mixing protocol, using a multi-inlet vortex reactor (MIVR) (Figure 5).92 The latter device has already
been evaluated for the large-scale production of a multitude of NPs.84
Figure 5. Graphical representation of a multi-inlet vortex reactor (MIVR) as a continuous mixing
geometry for polymer nanoprecipitation. (A) In an MIVR geometry, the highest energy mixing of
organic (solvent) and aqueous (anti-solvent) phase is observed at the outlet of the mixing chamber.
Adapted from ref
84
with permission from Elsevier, copyright 2011 (B) Schematic drawing of drug-
loaded polymeric nanoparticles coated with a PEGylated lipid monolayer, as typically obtained via a
nanoprecipitation protocol. Redrawn from ref 92.
A typical MIVR geometry is composed of two or four radially symmetric inlets converging into a
cylindrical mixing chamber. Separate inlets organize a continuous flow of organic phase (containing
the dissolved hydrophobic polymers) and antisolvent aqueous phase (containing the dispersed lipids)
16
into the central mixing area, initiating instantaneous nanoprecipitation and self-assembly of
homogenous lipid-coated polymer NPs.92 Employing this efficient and rapid mixing method, Fang
and coworkers achieved NP production rates > 10g/h, again without compromising the
physicochemical characteristics relative to equivalent hybrid NPs obtained via the previously
mentioned lab-scale sonication method.92 Alternatively, microfluidic-based approaches have been
gaining momentum in recent years for the well-controlled, reproducible and large-scale
manufacturing of polymer NPs via self-assembly.93 Much of the incentive in this context originated
from the engineering of microfluidic designs that enable rapid mixing of solvent and antisolvent to
obtain small and monodisperse polymeric NPs.93-95 Microfluidic channels can be engineered to
encompass internal micromixing structures that enable rapid mixing of focused liquid streams
flowing through these microchannels. Valencia and coworkers applied a so-called Tesla micromixing
structure to regulate the nanoprecipitation of lipid-coated PLGA NPs.96 In analogy with the MIVR
mixing geometry described above, separate inlet streams hydrodynamically focus a flow of organic
solvent and aqueous antisolvent through a microchannel with internal Tesla structure to induce rapid
homogenous mixing, resulting in the formation of (PEGylated) lipid-PLGA nanospheres with high
reproducibility.96 Aiming to optimize the high-throughput synthesis of these lipid-polymer hybrids, a
three-inlet three-dimensional (3D) microfluidic platform was designed that generates two symmetric
microvortices at the inlet intersection, controlling rapid mixing of PLGA polymer and lipid (Figure 6).
Interestingly, this platform demonstrated a productivity up to 3g/h at optimized flow rate and PLGA
concentration, significantly outperforming production rates obtained with the Tesla mixing design.97
Altogether, the compelling technological advances described here appear very promising toward
mass production of complex lipid-polymer nanocomposites with acceptable reproducibility and may
stimulate to a great extent their clinical translation.
17
Figure 6. Microfluidic-based nanoprecipitation method allowing the fast and reproducible synthesis
of PEGylated lipid-enveloped PLGA nanoparticles with low polydispersity. (A and B) Schematic
representation of a three-inlet microfluidic platform yielding microvortices for rapid yet controlled
mixing of PLGA polymer and lipid. (C) Structure of the obtained hybrid nanoparticles. (D) Microvortex
mixing demonstrates substantially improved production rates when compared with other
preparation methods. (Reprinted with permission from ref
97
. Copyright 2012 American Chemical
Society).
3.2. Lipid-coated mesoporous silica nanoparticles
The use of inorganic NPs, e.g. constructed from silica, for drug delivery purposes has increased
substantially the last decades.67,98 Following the successful deposition of (phospho)lipid bilayers onto
planar silica substrates 99 and µm-sized silica beads,21 many research groups initiated the decoration
of silica nanoparticles with a so-called supported lipid bilayer (SLB).21,100-102 Mornet and co-workers
described the successive steps in the adsorption of small unilamellar vesicles (SUVs) of varying
composition onto preformed silica nanoparticles via cryotransmission electron microscopy (cryoTEM). SUV adsorption, deformation and rupture at the lipid-particle interface precede the formation
of bilayer patches that subsequently coalesce to form a continuous lipid bilayer shell around the
anionic silica template. It was clearly visualized that the SLB closely followed the contours of the silica
18
core. Electrostatic interaction facilitated this process as judged by the superior SLB formation using
cationic or zwitterionic lipids opposed to lipid mixtures with a high net negative charge.100 In the
latter report solid silica NPs were produced according to the well-known Stöber process. However,
from a drug delivery perspective mesoporous silica nanoparticles (MSNPs) are clearly preferred.103
Generally, MSNPs are produced via a modified silica condensation protocol, in which supramolecular
assemblies of (cationic) surfactants above their critical micelle concentration (CMC) (i.e. spherical
micelles and liquid crystalline mesophases) are applied as soft templates during the sol-gel
conversion. Following condensation of the silica precursor around the surfactant head groups and
surfactant removal via calcination or solvent extraction, silica NPs are formed with a well-defined
uniform internal porous structure.104 Excellent reviews are available describing the various
preparation protocols of MSNPs in more detail.103,105-107 MSNPs are endowed with a tailorable pore
size (1.5 to several tens of nanometers), high specific surface area and large internal pore volume
which provide obvious assets toward encapsulation of drugs and imaging agents.106 A broad
spectrum of drugs can be retained in the mesopores as a function of pore size and pore structure as
well as direct physicochemical interaction between the guest molecules and the (modified) MSNP
surface.103,105-107 Consistent with early reports on solid silica NPs, successful lipid modification of
MSNPs has been demonstrated extensively.57,108-113 Sealing the MSNP pores by virtue of a SLB could
provide added protection of the precious drug cargo against the harsh in vivo microenvironment and
assist in preventing premature drug release in the extracellular matrix. The latter is illustrated by the
fast release of a protein drug payload at physiological pH from bare MSNPs (reaching 100% release
within 12h) while the lipid-coated MSNPs did not release relevant amounts of the protein as
measured over several days.113 The group of Jeffrey Brinker demonstrated the cargo loading and
concomitant pore sealing of conventional negatively charged MSNPs and cationic MSNPs modified
with 3-[2-(2-aminoethylamino)ethylamino]-propyltrimethoxysilane (AEPTMS) through fusion of
oppositely charged liposomes on the MSNP surface, to obtain so-called protocells. Again,
electrostatic interaction appeared to be the predominant factor driving SLB formation.108,110
19
Unfortunately, cationic lipids are notorious for their potential in vitro and in vivo toxicity (see also
section 5)
114-116
and charged nanomaterials are generally more prone to nonspecific
interactions.6,39,41 To reduce nonspecific binding and cellular toxicity, both anionic and cationic
MSNPs were incubated with liposomes consisting of zwitterionic lipids, cholesterol and PEGylated
lipids (Figure 7).57 In an alternative approach, Wang and coworkers modified the surface of
preformed MSNPs with long hydrophobic tails through reaction with 13-(chlorodimethylsilylmethyl)heptacosane. The hydrophobized MSNPs were subsequently capped with (PEGylated) phospholipids
in an organic solvent mixture. Following solvent evaporation and rehydration in salt-rich buffer,
phospholipid-coated MSNPs are formed via self-assembly of the hydrophobic heptacosyldimethylsilyl
groups grafted on the MSNP surface with the fatty acyl chains of the added phospholipids. The
resulting PEGylated phospholipid-capped MSNPs displayed superior colloidal stability and resistance
to nonspecific protein absorption as compared with unmodified MSNPs.109,117 Likewise, Koole et al.
functionalized silica-coated QDs with octadecanol, prior to their dispersion in a chloroform/methanol
mixture containing PEGylated lipids. Upon emulsification in an aqueous buffer and evaporation of
the organic solvent via a heating step, a monolayer of PEGylated lipids was deposited on the surface
of the hydrophobized silica NPs.35 Roggers and coworkers covalently modified MSNPs with a
dipalmitoyl monolayer to facilitate phospholipid coating using a comparable solvent evaporation
protocol.118 Organic solvents were also employed by Cauda et al. who incubated preformed drug
loaded MSNPs in an 40% (v/v) ethanolic lipid solution in order to obtain a defect-free lipid bilayer
coating. Upon dilution of the MSNP dispersion in water, decreasing the ethanol content to 5%, the
dissolved phospholipids self-assembled into an SLB shell around the MSNP core. This procedure was
successfully evaluated for both zwitterionic and cationic lipid mixtures.119
20
Figure 7. Schematic representation of lipid-enveloped mesoporous silica nanoparticles (termed
protocells). The nanoporous silica core serves as a matrix type carrier for the encapsulation of a
variety of therapeutic cargo and/or imaging agents. The silica core is surrounded by a supported lipid
bilayer (SLB), mainly comprising zwitterionic lipids in a fluid (1,2-dioleoyl-sn-glycero-3phosphocholine; DOPC) or more rigid gel-like phase (1,2-dipalmitoyl-sn-glycero-3-phosphocholine
DPPC). The SLB is further supplemented with 30 wt% cholesterol, 5 wt% PEGylated lipids and 1-5 wt%
of
1,2-dioleoyl-sn-glycero-3-phosphoethanolamine
(DOPE)
or
1,2-dipalmitoyl-sn-glycero-3-
phosphoethanolamine (DPPE) of which the primary amines are consumed for covalent anchoring of
targeting and fusogenic peptides via a heterobifunctional PEG spacer. (Reprinted by permission from
Macmillan Publishers Ltd: Nature Materials 57, copyright 2011).
3.3. Lipid-coated hydrogel nanoparticles
Hydrogels can be defined as 3D networks of hydrophilic polymers that are capable of absorbing large
quantities of water or biological fluid, while maintaining their internal network structure. Hydrogel
networks can be constructed by virtue of chemical and/or physical crosslinks between natural or
synthetic polymers. These crosslinks are the driving force behind the structural stability of the
21
hydrogel by preventing the dissolution of the hydrophilic polymers in the aqueous
microenvironment.120,121 Since their introduction for biological use more than 50 years ago,122 at
present hydrogels are favored for many biomedical and pharmaceutical purposes, including
controlled drug delivery.120,123-125 Hydrogels may exist in many geometries such as macroscopic gels
(e.g. scaffolds, cylinders,...), microscopic gels (microgels) and hydrogel nanoparticles (nanogels).
While the literature has predominantly focused on macro- and microscopic hydrogels for
extracellular drug delivery applications, during the last decade nanogels have gained momentum as
drug delivery vehicles.126 Nanogels maintain equal network properties compared to their micro-and
macroscopic counterparts, but their nanosized dimensions render them useful for intravenous
administration (e.g. for tumor targeting) and intracellular drug delivery. In addition, their small size
guarantees a rapid response to external stimuli, making them ideal candidates for triggered drug
delivery.126 To date, nanogels have mainly been employed as carriers for low molecular weight
chemotherapeutics, although in recent years the focus has shifted somewhat to macromolecular
drug delivery, including proteins and nucleic acids. Many recent review papers comprehensively
describe the plethora of available crosslinking methods, nanogel synthesis techniques and hydrogel
drug release mechanisms, which will not be covered here.126-128 Instead we will focus on nanogel-lipid
core-shell composites.
Extensive work was done by several groups on the lipid-coating of hydrogel microspheres for
controlled drug release or as an artificial cell mimic.21,129-136 The adsorption of a lipid bilayer onto the
microgel surface can be promoted by electrostatic interaction between charged microgels and
oppositely charged liposomes.130,132 Alternatively, introduction of hydrophobic anchors at the
microgel surface could drive the self-assembly of a (phospho)lipid bilayer in a subsequent
step.129,134,135 However, only a limited number of publications can be found that extrapolate these
findings to nanosized dimensions. Nevertheless, already more than two decades ago, a hybrid lipidnanogel drug delivery system was introduced, originally constructed to mimic low density
lipoproteins (LDL).137 The nanostructures, which were named SupraMolecular BioVectors (SMBV™),
22
consisted of a crosslinked natural polysaccharide core layered by a (protein modified) phospholipid
shell and were fabricated via a two-step synthesis. Preformed dextran nanogel cores were initially
acylated by grafting fatty acids of distinct chain length onto their surface. Acylated cores and
phospholipids were dispersed in ethanol and subsequently injected in an aqueous medium at
elevated temperature, resulting in the deposition of a phospholipid monolayer around the fatty acid
modified nanogel core.137 The SMBV concept was further extended to charged polysaccharide cores
via the introduction of quaternary ammonium or succinate/phosphate groups and at a later stage the
two-step synthesis was simplified by omitting the acylation step prior to phospholipid coating.54,138,139
Alternatively, dextran NPs have been coated with phospholipids via direct hydration of a dried lipid
film followed by high-energy sonication and extrusion.140
The majority of the synthesis methods used to prepare nanogels afford limited control over nanogel
size and/or size distribution and are often incompatible with labile macromolecular therapeutic
compounds such as pharmaceutical proteins.128 Protocols for lipid-coating of nanogels again mainly
consist of multiple consecutive steps that need separate optimization which makes the preparation
process laborious and difficult to control. Thus, single-step preparation protocols allowing
concomitant loading of drugs with disparate physicochemical properties and robust control over
nanoparticle dimensions, are highly sought after. Of particular interest in this regard, is the
exploitation of liposomes as nanoscaled reactors for the selective hydrogel formation in their
aqueous lumen. To this end, unilamellar liposomes are formed in the presence of a monomer
solution (e.g. via a thin-film hydration method accompanied by freeze-thawing, sonication or
extrusion), followed by a dilution-step and initiation of hydrogel crosslinking (Figure 8).
23
Figure 8. Liposomal-template assisted synthesis of hydrogel particles yielding lipid-coated nanogels.
Hydration of a thin lipid film with a (meth)acrylate functionalized monomer/polymer solution leads
to a dispersion of large multilamellar vesicles (LMVs) with broad size distribution. Extrusion of the
LMVs through a polycarbonate membrane with narrow pores is used to size the obtained
multilamellar vesicles into monodisperse large unilamellar vesicles or LUVs, carrying the
(meth)acrylated monomer/polymer in their aqueous core. Following dilution, radical polymerization
can be initiated, e.g. via UV illumination in the presence of a photoinitiator, resulting in the formation
of lipid-coated nanogels.
Dilution of the liposomal dispersion is of key importance to prevent macroscopic hydrogel formation
and thus to ensure selective monomer crosslinking in the liposome interior. Using this method, Van
Thienen et al. synthesized biocompatible PEG nanogels as well as biodegradable dextran nanogels
within a zwitterionic liposomal template.141-143 Kazakov et al. reported on the selective formation of
thermo-and pH-sensitive nanogels consisting of poly(N-isopropylacrylamide-co-1-vinylimidazole).
144,145
On the other hand, Schillemans et al. described the use of a detergent-dilution method to
construct the liposomal nanoreactor and relied on ascorbic acid to inhibit the free radical
polymerization process in between the liposomal vesicles.146 In another recent report, Park et al.
applied this liposomal template-assisted method in a protocol with sequential steps in which
preformed lyophilized liposomes were hydrated with a monomer solution.147 To initiate hydrogel
crosslinking, the latter reports made use of a water-soluble photoinitiator that was included in the
monomer solution prior to liposome formation, enabling UV-triggered photopolymerization to create
a chemically crosslinked hydrogel network. Alternatively, physically crosslinked alginate nanogels
were also successfully produced. Hereto, liposomes encapsulating sodium alginate in their aqueous
24
lumen were incubated in a solution of calcium chloride above the transition temperature of the lipid
bilayer. Permeation of divalent Ca2+ ions across the lipid bilayer shell could then initiate ionic gelation
of the negatively charged polysaccharide chains.148
3.4. Post-synthesis surface modifications
As briefly outlined in the introduction and comprehensively reviewed by others
39,41
, the surface
characteristics of a drug-loaded NP will impinge on how they interact with their biological
environment upon in vivo administration. Decorating the NP surface with a hydrophilic stealth layer
(e.g. using PEG) may safeguard colloidal stability and extend blood circulation. As indicated above,
for the construction of lipid-coated matrix NPs with a core-shell nanoarchitecture, this mostly implies
the use of PEGylated lipids during NP synthesis. However, some protocols may require the postsynthesis shielding with a PEG outer layer. Su et al. reported on the fabrication of pH-responsive
poly(-amino ester) NPs enveloped in a cationic PEGylated lipid shell.149 To this end, both a doubleemulsion solvent evaporation method and a solvent diffusion nanoprecipitation method were
investigated.
The
authors
found
that
adding
PEGylated
1,2-distearoyl-sn-glycero-3-
phosphoethanolamine (DSPE-PEG) during the nanoprecipitation protocol markedly reduced NP yield.
To resolve this issue, the DSPE-PEG was incorporated into the lipid bilayer via a post-insertion
process.149 Incubation of preformed conventional liposomes with a dispersion of PEG-modified lipids
above the critical micellar concentration (CMC) may indeed result in transfer of PEG-lipids from the
micellar phase into lipid bilayers in a time-and temperature dependent manner.150,151
Although of great importance, the presence of a hydrophilic polymer layer generally precludes
efficient endocytic uptake and processing by target cells due to reduced interaction with the cell
plasma membrane.30,152 Inserting PEGylated lipids with distal end functional groups into the NP
supported SLB, allows further post-synthesis modification of the NP surface with bioactive moieties
aiming to combine stealth properties with active cell targeting.2,3 For instance, thiol-maleimide
surface chemistry has been widely employed in this context by virtue of PEGylated lipids carrying
25
maleimide end-groups. In this way, lipid-coated PL(G)A NPs could be bestowed with cellular or tissue
targeting capability by covalently coupling cysteine-terminated peptides
aptamers
61
72
, thiolated 2’-OMe RNA
or disulfide reduced monoclonal antibody fragments.56 Likewise, Look et al. prepared
stealth liposomes using amine-terminated DSPE-PEG in order to covalently link nondepleting CD4
targeting antibodies using sulfo-NHS/EDC carbodiimide-based coupling chemistry. These
immunoliposomes were subsequently used for the preparation of so-called nanolipogels via the
liposomal-template assisted synthesis method (see also 4.3). Post-synthesis carbodiimide-assisted
modification was also applied by others to obtain lipid-coated PLGA NPs functionalized with the v 3
integrin-targeting Arg-Gly-Asp (RGD) peptide.58 Some lipids carrying less fragile targeting moieties,
such as DSPE-PEG-folate, can also readily be used during hybrid NP synthesis in the presence of
organic solvents.117,153-155
Functional moieties can also be covalently anchored directly to the SLB surface as demonstrated by
Brinker’s group for their silica based protocells (Figure 7). They coupled bioactive (fusogenic or
targeting) peptides, premodified for conjugation with a glycine-glycine spacer and a C-terminal
cysteine, to the primary amines of phosphoethanolamine (PE) polar headgroups via an amino-tosulfhydryl heterobifunctional succinimidyl-[(N-maleimido-propionamido)-tetracosa-ethyleneglycol]
ester (SMPEG24) crosslinker.57,112,113 The outstanding cellular targeting efficacy achieved with these
protocells clearly certifies that the presence of 5wt% DSPE-PEG, necessary to maintain colloidal
stability, in the silica supported SLB does not sterically interfere with ligand-receptor binding.57 Zheng
et al. reported on the modification of 1,2-dioleoyl-sn-glycero-3-phosphoethanolamine (DOPE), in
which the PE headgroup was reacted with a maleimidobenzoic acid N-hydroxysuccinimide ester (MBNHS), with thiolated transferrin (Tf). The obtained Tf-DOPE micelles were subsequently employed for
post-insertion into nanoprecipitated lipid-coated PLGA NPs.156
As detailed above, the majority of the NP production protocols encompass harsh preparation
including among others the use of organic solvents, the induction of shear stress via sonication, and
26
heating conditions, that might well be detrimental for labile drug molecules, e.g. nucleic acids and
proteins. For this reason, some authors prefer to add the therapeutic compound post-synthesis. This
strategy is frequently employed for the electrostatic complexation of anionic nucleic acids onto the
cationic surface of lipid-polymer hybrids.82,88,149,157 In this way, not only the possible degradation of
the therapeutic nucleic acid is avoided during synthesis, it may also remedy the low drug loading
often observed for these large hydrophilic compounds, e.g. in unmodified PLGA NPs.69,158,159 On the
other hand, it remains questionable to what extent this NP loading procedure can protect the nucleic
acids against enzymatic degradation and preterm decomplexation in vivo. In contrast, Bershteyn et
al. used a covalent coupling strategy to link thiolated ovalbumin (OVA) to the surface of lipidenveloped PLGA NPs via the maleimide functional group of DSPE-PEG-mal. To this end, the
ovalbumin antigen was premodified using a heterobifunctional crosslinker N-succinimidyl-S-acetyl(thiotetraethyleneglycol) of which the succinimidyl moiety can react with OVA primary amines. The
sulfhydryl group was deacetylated with hydroxylamine prior to thiol-maleimide coupling.
Interestingly, in a final step lipophilic danger molecules were also post-inserted in the lipid shell of
antigen-conjugated hybrids.52
4. DRUG DELIVERY APPLICATIONS OF LIPID-POLYMER MATRIX NANOCOMPOSITES
4.1. Lipid-coated polyester nanoparticles in drug delivery
4.1.1. Targeted delivery of chemotherapeutics
As described in section 3.1, Chan et al. optimized a single-step nanoprecipitation method for the
development of PLGA-lecithin-PEG core-shell NPs comprising a hydrophobic PLGA core, a soybean
lecithin monolayer and a non-fouling hydrophilic PEG outer shell.87 This NP platform (termed
‘nanoburrs’) was subsequently evaluated in vivo for targeted drug delivery to sites of injured
vasculature. To this end, paclitaxel was conjugated to PLA to obtain the hydrophobic drug-eluting
27
core. A heptameric KLWVLPK peptide, acquired via a phage display screen against collagen IV, was
conjugated to the PEG distal end via maleimide-thiol chemistry (Figure 9).
Figure 9. Overview synthesis of lipid-PLGA core-shell ‘Nanoburr’ particles. (A) synthesis scheme of
paclitaxel (Ptxl)-poly(lactic acid) (PLA) conjugate. (B) RP-HPLC on Ptxl and Ptxl-PLA conjugate. (C)
Illustration of hybrid nanoparticle formation via a single-step nanoprecipitation method. The
nanoparticles were modified post-synthesis with targeting peptides via maleimide-thiol chemistry.
(D-F) Morphological and physicochemical characterization of the Nanoburrs as a function of peptide
conjugation. (G) In vitro release profile of Ptxl obtained with Nanoburr NPs. (Reprinted with
permission from ref 72. Copyright 2010).
Targeted NPs displayed markedly enhanced accumulation at sites of compromised vasculature
following intraarterial and intravenous administration in an in vivo rat model of angioplasty-induced
vascular damage. The latter procedure may indeed disrupt the endothelial cell lining, thereby
exposing the underlying basal lamina enriched in collagen IV to which the paclitaxel-loaded NPs can
selectively bind.72 Systemic post-operative administration of targeted paclitaxel NPs achieved
significant reduction in neointimal smooth muscle cell proliferation and vessel restenosis, as
28
compared to the standard micellar paclitaxel formulation (Taxol) and non-targeted controls.160 This
targeted drug delivery approach could also prove beneficial toward treatment of pathologies
accompanied by increased vascular permeability, such as cancer and cardiovascular inflammatory
diseases, allowing the drug-loaded particles to locally extravasate and bind the collagen-rich
basement membrane, combining passive EPR-mediated accumulation with active ligand-mediated
targeting.72 The same type of PEGylated lipid-PLGA NP has also been evaluated for targeting
approaches focused on reaching cancer cells instead of compromised vasculature. The group of
Liangfang Zhang modified the lipid monolayer surface with selectively reduced half-antibody
fragments targeting the carcinoembryonic antigen (CEA) that is frequently overexpressed in
pancreatic cancer. The anti-CEA antibody-conjugated NPs displayed superior targeting specificity
towards CEA expressing cancer cells that translated in an enhanced cytotoxic effect of paclitaxel
compared with non-targeted control.56
4.1.2. Combination drug delivery
The controlled co-delivery of multiple drugs via a single nanocarrier has the potential to overcome
the drawbacks that are frequently associated with monotherapy.161 In addition, ample control over
the drug delivery process in a spatiotemporal fashion may entail a synergistic therapeutic effect. One
of the key challenges in this regard is the efficient formulation of drug molecules with varying
physicochemical properties in the same nanocarrier in a controllable and reproducible manner. To
this end, Aryal et al. reported on the synthesis of amphiphilic drug conjugates of two widely used
chemotherapeutic agents, i.e. the hydrophobic paclitaxel and hydrophilic gemcitabine hydrochloride,
which were subsequently encapsulated into lipid-coated PLGA NPs via a single-step
nanoprecipitation protocol (Figure 10).
29
Figure 10. Synthesis scheme of the hydrolysable paclitaxel-gemcitabine conjugate. (Reprinted with
permission from ref 162. Copyright  2010 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim).
The drug conjugate loaded NPs demonstrated markedly enhanced cytotoxicity compared to free
conjugate in a human pancreatic carcinoma cell line. The intracellular combinatorial release of both
chemotherapeutics is believed to occur via PLGA degradation and concurrent acid-catalyzed
hydrolysis of the drug conjugate ester linkage in acidified endolysosomal vesicles.162 In a similar
report, the same authors also demonstrated enhanced ovarian cell killing by encapsulating a
hydrolysable paclitaxel-cisplatin drug conjugate in lipid-polymer hybrid NPs.163 However, no details
are provided on the functional role of the lipid monolayer in the cellular drug delivery process.
Sengupta et al. on the other hand elegantly demonstrated the benefit of a lipid-polymer core-shell
nanoarchitecture to gain better spatiotemporal control over the delivery of two distinct anticancer
drugs.73 These authors proposed PLGA NPs that covalently encapsulate doxorubicin allowing a slow
and degradation-controlled release of the chemotherapeutic. In addition, the PLGA NPs were
enclosed in a PEGylated lipid bilayer in which the hydrophobic anti-angiogenesis drug combretastatin
was inserted. Following systemic administration, the PEGylated lipid shell ensures prolonged blood
circulation times and increased extravasation in the tumor interstitium by virtue of the EPR effect.
The physical segregation of both therapeutic agents allowed a temporal drug release profile with a
fast initial release of the combretastatin, causing angiogenesis inhibition, preceding a slow release of
the cytotoxic doxorubicin. This sequential exposure of the tumor to two distinct types of drugs
synergistically acting on both the cancerous cells and the tumor vasculature allows the accumulation
30
of effective doxorubicin concentrations in the tumor mass.73 Of note, recent insights demonstrated
hindered extravasation of larger nanomedicines (>100 nm) following tumor vessel normalization by
blocking vascular endothelial growth factor (VEGF) receptor-2.164 This report thereby emphasizes the
importance of NP size in this therapeutic context, especially when repeated administration is
required, given that only the smallest NPs (~12 nm) showed improved tumor penetration following
VEGF treatment. In a comparable approach, Wang and Ho formulated combretastatin together with
paclitaxel in RGD-modified lipid-PLGA nanohybrids.58 The RGD peptide, conjugated to the distal end
of PEGylated lipids inserted in the lipid bilayer, is targeted to v 3 integrin receptors highly expressed
on neovascular blood vessels in tumors 165. In this report, the lipid bilayer serves both to facilitate the
encapsulation of combretastin as to sustain the release of the paclitaxel from the PLGA core.58
In
another
combinatorial
strategy,
Wang
and
coworkers
translated
the
concept
of
chemoradiotherapy, i.e. the synchronized treatment with both chemotherapy and radiotherapy, to
lipid-PLGA nanomedicines. The utilization of a lipid-polymer hybrid NP platform
60
for
chemoradiation, which they fittingly dubbed ChemoRad NPs, enabled the encapsulation of radioisotopes without affecting the NP surface characteristics or interfering with drug encapsulation
efficiency and drug release kinetics.59 The latter was accomplished by the integration of
phospholipids, modified with the chelator diethylenetriaminepentaacetate (DTPA), in the PEGylated
lipid monolayer that surrounds the docetaxel loaded PLGA core. By employing the A10 RNA aptamer
as a targeting ligand, ChemoRad NPs demonstrated selective delivery of docetaxel and yttrium90 (90Y)
to prostate-specific membrane antigen (PSMA) overexpressing prostate cancer cells and showed
synergistic cell killing when compared with targeted monotherapy.59 Likewise, folate-targeted
ChemoRad NPs were engineered by others for the concurrent delivery of paclitaxel and 90Y. Following
their initial identification of folate-targeted and docetaxel-loaded hybrid NPs as effective
radiosensitizers
154
, the authors aimed to evaluate co-delivery of chemo-and radiotherapeutics in a
mouse model of ovarian cancer peritoneal metastasis. Intraperitoneal injection of targeted
ChemoRad NPs significantly outperformed their non-targeted counterparts and combinatorial
31
chemoradiotherapy achieved the most effective therapeutic outcome.153 Folate-targeted, PEGylated
and DTPA modified lipid-PLGA NPs were also proposed as targeted nanotheranostic agents, in which
the DTPA chelator accommodates gadolinium (Gd3+) as an MRI-contrast agent.83
PLGA NPs have also been employed for the simultaneous loading of chemotherapeutics and thermooptical agents.166 Within the latter group, the FDA-approved near-infrared (NIR) fluorescent dye
indocyanin green (ICG) has been frequently applied in photodynamic and photothermal
therapy.166,167 Upon absorption of NIR photons, ICG has demonstrated to emit the excitation energy
as heat, a trait that can be exploited to induce local hyperthermia.166 This localized heating effect is
believed to synergize with standard chemotherapy by sensitizing cancer cells to its cytotoxic
activity.168 Aiming to combine both chemotherapy and photothermal therapy in a single nanoparticle
delivery system, Zheng and coworkers recently developed PLGA-lecithin-PEG NPs encapsulating a
mixture of doxorubicin and ICG employing a single-step sonication method. They found that the
combined chemophototherapy entailed a synergistic tumor cell apoptosis (Figure 11) and tumor
growth inhibition in a mouse MCF-7 xenograft model.169 Unfortunately, the in vivo administration of
the lipid-coated PLGA NPs was performed via intratumoral administration, thus bypassing much of
the pre-existing barriers associated with systemic intravenous injection.
32
Figure 11. Intratumoral injection of lipid-PLGA nanohybrids loaded with doxorubicin and ICG (DINPs)
in nude mice MCF-7 breast cancer xenografts. (A-B) Laser irradiation causes local hyperthermia due
to ICG. (C) Thermally induced tissue damage and (D) tumor cell apoptosis (green staining, TUNEL
assay) following laser irradiation in the presence of free ICG, ICG-loaded NPs (INPs) or DINPs. (E)
Chemo-photothermal combination treatment with DINPs caused a marked anticancer effect,
outperforming free drug and INP monotherapy. (Adapted with permission from ref
169
. Copyright
2013 American Chemical Society).
4.1.3. Triggered drug delivery
A desirable trait of nanomedicines in general is the ability to respond to an external or biological
environmental stimulus (pH, light, temperature, magnetic field,…) with the aim to achieve better
33
spatiotemporal control over the drug delivery process, thereby enhancing therapeutic efficacy at the
biophase and minimizing off-target toxicity.170 Both liposomes and polymeric NPs have been explored
extensively as stimuli-responsive nanocarriers, paving the way toward more complex hybrid
nanocomposites for triggered drug delivery approaches.171 More specifically in the context of lipidpolymer nanocomposites, magnetic Fe3O4 NPs, incorporated in lipid-coated PLGA hybrids together
with the anticancer drug camptothecin, have been shown to dictate the release of the latter upon
applying an external magnetic field. The lipid shell has the particular task of avoiding premature drug
release in absence of the external stimulus.172 Clawson and coworkers synthesized a PEGylated lipid
with a pH-sensitive succinate merging the dipalmitoyl phospholipid and the PEG2000 moiety.173 Lipidpolymer nanohybrids were constructed with a PLGA core and a fusogenic lipid monolayer consisting
of DOPE and oleic acid. The interspersed PEGylated lipids ensure steric stabilization of the NPs a
neutral pH, yet drive lipid fusion and NP aggregation upon succinate hydrolysis and PEG shedding at
acidic pH. The NP destabilization as a function of pH and incubation time can be tuned by adjusting
the molar ratio of pH-sensitive PEGylated lipids incorporated in the lipid monolayer. It is anticipated
that this pH-dependent behavior would be of particular interest in cancer drug delivery by taking
advantage of the acidified tumor interstitium.174 Selective removal of the stabilizing PEG layer in the
tumor microenvironment can thus induce fusion with the tumor cell membrane. In addition, gradual
acidification of endosomal compartments following endocytosis of PEGylated NPs in tumor cells, can
trigger lipid fusion with the limiting membrane of late endosomes or lysosomes. Both mechanisms
could greatly enhance cytosolic drug delivery.173
4.1.4. Delivery of macromolecules
Next to low molecular weight chemotherapeutics, lipid-coated PLGA NPs have also been assessed
with respect to delivery of macromolecular drugs, including proteins and nucleic acids. The group of
Darrell Irvine applied PLGA micro-and nanoparticles for protein antigen delivery in a prime-boost
vaccination regimen. As further detailed in section 4.4., in this report the antigens were not
34
dispersed in the PLGA matrix core yet anchored onto the lipid surface, inspired by natural pathogenic
antigen presentation.52 In recent years, genetic vaccination, making use of plasmid DNA (pDNA) or
messenger RNA (mRNA) to drive protein antigen expression and presentation by dendritic cells (DCs),
has gained momentum. In this context, mRNA vaccination has important advantages over DNA-based
strategies. In contrast to pDNA, mRNA cannot insert into the host genome, excluding the risk of
insertional mutagenesis. Moreover, mRNA does not require delivery into the nucleoplasm, thereby
avoiding the barrier imposed by the nuclear envelope and allowing transfection of non-dividing
cells.175 Su and coworkers designed fully degradable and pH-responsive poly(-amino ester) (PBAE)
NPs for mRNA vaccination. To minimize cytotoxicity of the hydrophobic polycation core, a PEGylated
lipid envelope was created around the NPs. Additionally, to avoid exposure of the labile mRNA to the
harsh conditions during NP preparation, 20 mol% of the cationic lipid 1,2-dioleoyl-3trimethylammonium-propane (DOTAP) was included in the lipid mixture, thus allowing the postsynthesis complexation of mRNA to the positively charged lipid surface. It was demonstrated that the
presence of a PEG corona did not interfere with mRNA loading. Successful mRNA delivery with these
particles was demonstrated in vitro in the murine dendritic cell line DC2.4 and in vivo following
intranasal administration.149 Next to mRNA, the use of cationic lipids during NP preparation to
enhance (surface) complexation of pDNA 82 and siRNA 88,157 has been demonstrated by others. Folate
targeted lipid-PLGA NPs were recently proposed for the co-delivery of doxorubicin and pDNA, which
were encapsulated in the hydrophobic PLGA core or electrostatically complexed to the cationic outer
lipid shell, respectively.82 Yang et al. prepared PEGylated PLA NPs with a cationic lipid monolayer at
the interface of PLA core and PEG shell. Surface complexation of siRNA targeting the oncogene pololike kinase 1 (Plk1), significantly suppressed tumor growth in a murine xenograft model following
systemic administration (Figure 12).88
35
Figure 12. Formation of lipid-polymer hybrid nanoparticles via a single-step nanoprecipitation of
poly(lactic acid) (PLA) and the amphiphilic PEG-PLA conjugate in the presence of the cationic lipid
N,N-bis(2-hydroxyethyl)-N-methyl-N-(2-cholesteryloxycarbonyl aminoethyl) ammonium bromide
(BHEM-Chol). The cationic lipid monolayer at the interface of the PLA hydrophobic core and the PEG
hydrophilic shell allows surface complexation of siRNA targeting the polo-like kinase 1 (Plk1)
oncogene. RNAi silencing of Plk1 in vivo following systemic administration significantly suppressed
tumor growth (Adapted with permission from ref 88. Copyright 2012 American Chemical Society).
Aiming for topical siRNA delivery to treat inflammatory skin disorders, Desai et al. recently described
co-delivery of anti-TNF siRNA and capsaicin via formulation in PLGA NPs (encapsulating the
hydrophobic capsaicin), enveloped within a lipid shell consisting of cationic amphiphiles carrying
cyclic pyrrolidinium head groups. The latter enables efficient complexation of the negatively charged
siRNA.157 Several strategies have been described in the literature to stimulate siRNA loading in PL(G)A
36
matrices by precomplexation with cationic polymers.176-178 Interestingly, Yang and coworkers applied
a proprietary cationic lipid to drive siRNA encapsulation in PLGA matrices during a modified double
emulsion solvent evaporation method using PVA as a final stabilizer. In the course of the primary
emulsification step, the cationic lipids tightly self-assemble at the water-oil interface, stabilizing the
aqueous droplets in which the siRNA will be complexed to the cationic lipid headgroups. In this way,
encapsulation efficiencies exceeding 90% are attained and the drug loading weight ratio almost
reached 5%.159 Shi et al. extended this method to obtain differentially charged lipid-polymer-lipid
nanostructures with hollow core for siRNA encapsulation due to the addition of a mixture of DSPEPEG and lecithin acting as stabilizers during the second emulsion step and solvent evaporation
(Figure 13). These hybrid lipid-polymer NPs were evaluated for siRNA delivery in vivo in a murine
subcutaneous xenograft model and could evoke moderate luciferase silencing upon a single
intratumoral injection.158
Figure 13. Lipid-PLGA-lipid hybrid nanoparticles with core-shell morphology designed for siRNA
delivery. With the aid of positively charged lipids, a high encapsulation efficiency of the anionic siRNA
is achieved in the nanoparticle hollow core. (A) Schematic representation and (B) transmission
electron microscopy (TEM) image of the hybrid nanoparticle construct. (C) Confocal image of hybrid
microparticles showing the outer PEG-lipid monolayer in green, the inner cationic lipid monolayer in
red and the in-between PLGA layer in blue. (Reprinted with permission from ref 158. Copyright  2011
WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim).
37
In recent years, it has been documented that besides NP size and surface chemistry, also NP shape
can have a profound effect on biodistribution, particokinetics and cellular internalization and that
nanospheres are not by definition superior over other NP design.39,76,179,180 Recognizing the
importance of nanogeometry on NP interactions at the nano-bio interface, monodisperse needleshaped PLGA particles (with dimensions of 80x320 nm) were fabricated by virtue of the PRINT
process introduced in section 3.1.. It has indeed been demonstrated that rod-like NPs with high
aspect ratio are more efficiently internalized by cells.179 The non-spherical PLGA NPs were utilized for
siRNA encapsulation without the aid of cationic lipids or polymers. In this way, an siRNA
encapsulation efficiency nearly reaching 50% was achieved. Nevertheless, to assist in cellular
internalization, a cationic lipid coat was layered around the particles post-synthesis. Three different
prostate cancer cell lines were successfully transfected with siRNA targeting the KIF11 (Eg5) gene
which encodes a kinesin-like motor protein that is of essential importance during cell mitosis.
Selective downregulation of this protein caused a marked decrease in cell viability in all three cell
lines tested. Regrettably, no in vivo data are available yet for this drug delivery platform.75 However,
it is conceivable that additional optimization, e.g. by bestowing the particles with targeting and/or
stealth properties, will be indispensable for successful in vivo translation.
4.2. Lipid-coated mesoporous silica nanoparticles in drug delivery
The multifunctional character of lipid-polymer nanocomposites was impressively illustrated by
Brinker’s group that reported on drug-loaded MSNPs functioning as nanoscopic templates for the
deposition of a phospholipid SLB.57 The resulting hybrid NPs were ultimately termed ‘protocells’,
referring to their elementary resemblance of a cellular construct. The nanoporous silica core
enabled encapsulation of high levels of various therapeutic (e.g. dsRNA, doxorubicin, diphtheria toxin
A-chain,…) and diagnostic agents (e.g. quantum dots) with distinct physicochemical properties. Next,
38
the drug-loaded core was sealed off with a single lipid bilayer mainly composed of zwitterionic and
PEGylated lipids with the primary aim to protect the therapeutic cargo from premature release and
reduce aspecific binding to nontarget cells. To facilitate specific cell targeting and intracellular drug
delivery, the lipid shell was further functionalized with a targeting peptide (SP94, targeting human
hepatocellular carcinoma (HCC)) and a pH-responsive peptide (stimulating disruption of both
endosomal membrane and SLB at endolysosomal pH) respectively (Figure 7). Importantly, by
wrapping a fluid-phase SLB around the MSNP core, the lipid-anchored targeting peptides retain their
lateral mobility within the bilayer. It was demonstrated that this particular feature allowed
multivalent receptor-mediated binding of target cells with minimal surface densities of peptide
ligand. Combining exceptional target cell specificity and drug loading capacity with proficient
cytosolic delivery, a 106-fold improvement in cancer cell killing was achieved over conventional
liposomes using protocells carrying a cocktail of doxorubicin, 5-fluorouracil and cisplatin.57
Unfortunately, to date no proof-of-concept data on this targeted drug delivery platform in validated
animal models are available yet and thus the in vivo targeted drug delivery performance of protocells
still remains elusive.181 In fact, it has recently been disclosed that the deposition of a protein corona
upon instillation of targeted nanomedicines in a biological medium may prominently alter their
ultimate targeting efficacy,
182,183
a finding that warrants further in vivo evaluation of targeted
protocells, or any other targeted NP for that matter.
As stated above, next to encapsulation of low molecular weight chemotherapeutics, MSNPs can also
be tuned to support loading with high molecular weight biomacromolecules. To this end, MSNPs with
a bimodal pore morphology were designed by using a mixture of two different types of surfactants as
structure-directing agents. This particular templating strategy resulted in MSNPs containing large
surface-accessible pores (~10-30 nm) interconnected by smaller pores (~5 nm).184 The larger pore
diameter should allow the penetration and retention of more bulky therapeutic molecules such as
proteins and nucleic acids in the MSNP matrix. Targeted protocells constructed with bimodal MSNP
cores as described above, were thus further pursued for the intracellular delivery of siRNA 112 and the
39
deglycosylated ricin toxin A-chain (RTA).113 In order to better accommodate the negatively charged
siRNA and RTA in MSNPs, the latter were made cationic by incorporating the amine-containing silane
AEPTMS, which significantly augments the MSNP loading capacity for both therapeutics. Moreover,
AEPTMS modified MSNPs show faster dissolution kinetics at physiological conditions, promoting the
release of encapsulated cargo.57,112,113 Ashley et al. encapsulated an equimolar mixture of distinct
siRNAs in the protocell core, designed to silence different members of the cyclin protein
superfamily.112 Cyclins are key regulatory proteins in various stages of the cell cycle that sustain
proliferation of malignant cells through activation of cell-cycle dependent kinases.185 Targeted
protocells could achieve maximal cyclin knockdown in the HCC cell line Hep3B with siRNA
concentrations in the low pM range and thereby significantly outperform standard cationic
liposomes. Importantly, the protocell-induced RNAi effect could evoke growth arrest and apoptosis
in HCC cells at a particle:cell ratio ~10, without affecting the viability of normal control hepatocytes
lacking expression of the antigen that is recognized by the protocell-anchored targeting peptide.112
Likewise, Epler et al. demonstrated the selective induction of apoptosis in target HCC cells by
cytosolic delivery of the protein toxin RTA.113 The catalytic ricin toxin A subunit inhibits protein
synthesis by cleaving a specific N-glycosidic bond in 28S ribosomal rRNA, thereby irreversibly blocking
mRNA translation.186 Protocells could achieve a half-maximal inhibition of protein synthesis at a RTA
concentration ~5pM, which was a 100-fold more potent than RTA loaded control liposomes with an
equal lipid composition as the protocell’s SLB (Figure 14).113
40
Figure 14. Protocell-guided toxin delivery. (A) Mesoporous silica nanoparticles (MSNPs), modified
with an amine-containing silane (AEPTMS) were loaded with the deglycosylated ricin toxin A chain
(RTA), capped with a fluid-phase PEGylated phospolipid bilayer and modified with both a targeting
(SP94) and an endosomolytic peptide, in line with reports by Ashley et al.
57,112
(B) Transmission
electron microscopy (TEM) image (scale bar = 50 nm) showing a MSNP core with bimodal porosity.184
The inset graph (scale bar = 200 nm) shows a scanning electron microscopy (SEM) image of a
microscopic MSNP to visualize the surface-accessible nanopores. (C) Inhibition of protein synthesis
induced by targeted RTA delivery in a hepatocellular carcinoma (HCC) cell line. (Reproduced with
permission from ref 113. Copyright  2012 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim).
The protocells described above are particularly well designed toward intracellular drug delivery as
they respond to the gradual acidification in the endosomal lumen following receptor-mediated
endocytosis and subsequent endosomal maturation.63 Endosomal pH likely reduces electrostatic and
dipolar interactions between the cargo loaded MSNP core and the PE/PC lipid polar headgroups,
resulting in SLB destabilization. In addition, protonation of the lipid-anchored fusogenic peptide
disrupts the endosomal membrane enabling cytosolic drug release.57,112,113 Besides variations in pH,
41
the intracellular reductive environment can also be exploited to control drug delivery.63 To this end,
thiolated MSNPs were decorated with a lipid bilayer of which the inner leaflet was covalently
attached to the MSNP surface via disulfide bonds. Triggered release of a fluorescent tracer molecule
was demonstrated upon disulfide reduction and lipid bilayer shedding.118 This drug delivery concept
however requires further intracellular confirmation with a model drug. Importantly, drug loaded NPs
containing disulfide bonds will preferentially disassemble in the reductive cytoplasm rather than in
endocytic compartments, again highlighting the importance of endosomal escape.6
Schlosbauer and coworkers recently reported on an alternative external light-triggered approach,
i.e. photochemical internalization (PCI), to overcome the endosomal barrier.187 PCI is based on the
selective accumulation of amphiphilic photosensitizers (PS) in the limiting bilayer of endocytic
vesicles. Excitation of the PS compound with light of the appropriate wavelength is followed by its
reaction with oxygen and induces the formation of reactive oxygen species (ROS), primarily singlet
oxygen (1O2). This highly reactive intermediate can cause oxidative damage to cellular components,
but this effect is mainly confined to the local production site of the singlet oxygen, owing to its short
range of action and short lifetime. This localized effect will therefore selectively disrupt the
endosomal membranes, releasing the endocytosed macromolecules or NPs into the cytosol.188,189
Since its discovery in 1999,190 PCI has been successfully applied to stimulate the cytosolic delivery of
several types of macromolecules (peptides, proteins, nucleic acids), incorporated in non-viral carrier
systems.191-193 However, one of the drawbacks of this technique on a cellular level is the inherent
cytotoxicity associated with the localization of PS molecules in the plasma membrane and other
organelles before illumination. To avoid this limitation, Schlossbauer et al. proposed to covalently
couple the PS to the surface of drug loaded MSNPs prior to their encapsulation in a SLB. In this way,
the PCI effect can be confined to endolysosomes that contain one or more lipid-coated MSNPs in
their lumen following cellular internalization. The authors could show that upon illumination of the
PS, protoporphyrin IX, the formed 1O2 sequentially disrupts both the SLB and the endosomal bilayer.
It was demonstrated in a human hepatoma cell line that this two-step phototoxic effect could result
42
in the successful cytosolic release of distinct membrane-impermeable drugs.187 On the other hand,
Teng et al. recently reported on protoporphyrin IX loaded, folate-targeted and phospholipidfunctionalized MSNPs for anti-cancer photodynamic therapy, exploiting the light-triggered ROS
induction to selectively induce apoptosis in cancer cells.117
The application of an external magnetic field can also be employed as a physical stimulus to promote
drug delivery.170 To this end, superparamagnetic iron oxide nanocrystals (SPIONs, ~20 nm) have been
encapsulated inside MSNPs (~100 nm) before capping them with a zwitterionic DOPC lipid bilayer.
Subjecting the particles to an alternating magnetic field causes local heating which affects SLB
permeability and triggers drug release.194 Although the SPION-MSNP hybrids seem to be reasonably
well tolerated by various human cell lines, the usefulness of these stimuli-responsive NPs toward
intracellular drug delivery still requires experimental validation. Alternatively, SPIONs have been
widely investigated as MRI contrast agents, possibly rendering such particles also valuable toward
theranostic applications.195
4.3. Lipid-coated hydrogel nanoparticles in drug delivery
Pioneering research on the lipid-coating of microscopic gel particles, e.g. performed by the groups of
Needham and De Smedt, clearly illustrate the potential of combining stimuli-responsive hydrogel
particles with a lipid bilayer coating for triggered drug delivery.130,132,133,196 Kiser and Needham
designed anionic pH-responsive poly(methacrylic acid) microgels for doxorubicin loading and
decorated these microgels with a phospholipid shell to control the drug delivery process. The
microgels were loaded with doxorubicin in a swollen state at physiological pH, exploiting the anionic
character of the gel to electrostatically bind the cationic doxorubicin. Decreasing the pH of the
medium condensed the drug loaded microgels due to the protonation of the methacrylic acid groups.
The condensed microgels were subsequently capped with a phospholipid shell that served as a
diffusion barrier to avoid drug leakage and to stabilize the microgel in its condensed state upon reimmersion in a pH neutral buffer. Permeabilization of the SLB via electroporation or lipid
43
destabilizing surfactants instigated pH-dependent swelling of the microgel core, disruption of the
lipid bilayer and release of doxorubicin.132,133 In another example, De Geest et al. coated
biodegradable anionic dextran microgels with an oppositely charged phospholipid layer and
demonstrated that the increased swelling pressure, as a function of the degradation of the microgel
core, could rupture the surrounding lipid membrane.130,196 In the same group, Van Thienen et al.
documented the liposomal-template assisted synthesis of equivalent degradable dextran nanogels
for protein delivery.142 In contrast to the earlier observations on a microscopic scale, it was
demonstrated that the lipid coat of these nanosized core-shell particles stays layered around the
nanogel core during its degradation. Likely the resulting build-up of internal osmotic pressure is
insufficient to overcome the tensile strength of the lipid membrane.141 On the other hand, the
swelling pressure of the degrading core could increase the drug permeability of the surrounding lipid
layer, as it was shown that the release kinetics of model proteins could be tailored from days to
weeks dependent on the crosslink density of the gels and the presence of a lipid coat.142 The
SupraMolecular BioVectors (SMBV™) designed by Biovector Therapeutics, which also contain a
polysaccharide dextran/maltodextrin core, have been evaluated for delivery of various
therapeutics.19 For example, antigen-loaded SMBVs have been put forward as nanoparticulate
vaccins for intranasal administration, mimicking natural viral pathogens both in terms of size (60-80
nm) as in antigen presentation. The SMBVs generated mucosal and serosal immunity in (pre-)clinical
assessment as vaccination strategy against influenza A.54 Next to proteins, also antisense
oligonucleotides have been encapsulated in SMBVs.197 Lipid-enveloped dextran NPs were additionally
investigated as a nanotheranostic platform by Erten and coworkers. The dextran nanogel matrix,
loaded with doxorubicin, was constructed around an iron oxide core as MRI contrast agent. The iron
oxide-dextran nanogels were endowed with a PEGylated lipid coat to enhance in vivo
biocompatibility. Furthermore, to maximize the physical integrity of the lipid coat, a fraction of
acetylated DOPE lipids was incorporated to enable mild UV-triggered crosslinking of the SLB. The
theranostic NPs were investigated in a murine tumor xenograft model where an enhanced MRI
44
contrast, indicative of NP accumulation, could be observed in tumors through a dorsal skinfold
window chamber.140,195
Very recently, the group of Tarek Fahmy impressively showed the implementation of nanosized lipidcoated polymeric gels (termed ‘nanolipogels’) for the immunotherapeutic treatment of metastatic
melanoma (Figure 15).147 The success of cancer immunotherapy largely depends on the
immunosuppressive nature of the tumor microenvironment. Many tumors can develop tolerance via
specific resistance mechanisms leading to immune evasion and failure of immunotherapeutic
strategies. Novel immunotherapeutic approaches should therefore aim at improving tumor
immunogenicity in order to attain an effective anti-tumor immune response.198,199 The accumulation
of tolerogenic cytokines, such as transforming growth factor- (TGF-), in the tumor tissue interstitial
space is one of the hallmarks that mitigate anti-tumor immunity.198,200 Therapeutic strategies aimed
at blocking the function of these immunosuppressive factors should be able to reinforce T-cell
mediated tumor cell killing.198 Park et al. therefore envisioned the combinatorial controlled delivery
of a commercially available TGF- receptor-I inhibitor together with IL-2, an immunostimulating
cytokine that activates endogenous cytotoxic T cells. However, a major impediment in this codelivery strategy is the co-encapsulation of high molecular weight, water-soluble IL-2 and the small
hydrophobic TGF- receptor-I antagonist in a single nanocarrier. To enable a concomitant sustained
release of both molecules, the authors therefore designed nanolipogels using PEGylated liposomes
as a nanoreactor for the selective photopolymerization of methacrylated -cyclodextrin and the
terminally diacrylated poly(lactic acid-co-ethylene glycol-co-lactic acid) biodegradable macromer.147
Cyclodextrins have been frequently applied for the formulation of poorly water-soluble drugs by
virtue of their ability to form molecular inclusion complexes with hydrophobic molecules.11 This
particular feature is exploited here to stably encapsulate the TGF- receptor-I inhibitor in the
liposome interior together with the IL-2 cytokine that is sterically entrapped in the degradable
polymer gelmatrix. Controlled degradation of the latter enabled sustained delivery of both
therapeutics. In a murine model of metastatic melanoma, the combination therapy entailed superior
45
survival benefit over either monotherapies following systemic administration. It was demonstrated
that the synergistic anti-cancer effect was the result of enhanced infiltration of both activated
cytotoxic CD8+ T cells as well as natural killer (NK) cells.147
In a follow-up publication, the nanolipogel formulation was also evaluated for the sustained delivery
of the immunosuppressant mycophenolic acid (MPA) for the treatment of the autoimmune disease
systemic lupus erythematosus (SLE). Likewise, the co-polymerization of -cyclodextrins in the
biodegradable gel-like core enabled efficient loading of the hydrophobic MPA. Intraperitoneal
administration of MPA-loaded nanolipogels significantly extended survival in a murine lupus model in
contrast to free drug. Although the involvement of CD4+ T cells in SLE pathophysiology has been well
established, the formulation of MPA in CD4-targeted nanolipogels did not give rise to a clinical
benefit compared to their non-targeted counterparts. Biodistribution studies indicated that both
nanoformulations are preferentially captured by macrophages and dendritic cells in lymphoid organs.
Delivery of MPA to DCs resulted in decreased secretion of inflammatory cytokines, likely explaining
the in vivo therapeutic outcome.53
46
Figure 15. Schematic overview of the synthesis protocol of drug-loaded nanolipogels. (A)
Methacrylated -cyclodextrins are employed for the solubilization of the hydrophobic TGF- inhibitor
SB505124. (B) Preformed and lyophilized PEGylated liposomes are loaded with the methacrylated
cyclodextrin inclusion complex, a biodegradable diacrylated poly(lactic acid-co-ethylene glycol-colactic acid) polymer and the cytokine IL-2. The PEGylated liposomes next serve as nanosized
templates for the selective UV-induced photopolymerization of the (meth)acrylated compounds,
thus creating a hybrid core-shell nanolipogel particle. (Reprinted by permission from Macmillan
Publishers Ltd: Nature Materials 147. Copyright 2012).
47
4.4. Bio-inspired and bio-mimetic lipid-coated nanoparticles in drug delivery
Synthetic nanomedicines often fail to surmount the numerous biological barriers en route to their
(intracellular) drug target. Poor efficacy in delivering therapeutic concentrations of a drug into the
target tissue or cell and eminent (immune)toxicity frequently hamper the clinical translation of
fundamental nanomedicine concepts.201 Fuelled by these limitations, currently a growing interest
exists in the implementation of bio-inspired and bio-mimetic materials and drug delivery
approaches.202,203 Naturally occurring micro-and nanoparticulate systems (e.g. lipoproteins,204,205
extracellular vesicles,206 mammalian cells90) are believed to possess specific features, including
improved in vivo stability, biocompatibility and intrinsic cell/tissue targeting, from which drug
delivery carriers could benefit.
Recent progress in nanotechnologies and emerging knowledge on the properties of biological
particulates should jointly pave the way toward successful hybrid bio-mimetic nanocomposites.203 In
the context of core-shell lipid-polymer hybrid NPs, strategies are materialized to endow (synthetic)
polymeric NPs with a phospholipid bilayer of natural origin. Red blood cells (RBCs) have already been
extensively explored as drug delivery carriers owing to their biocompatibility and their extended in
vivo circulation time (~120 days).203,207 Inspired by these favorable characteristics, Hu et al. recently
reported on the coating of polymeric PLGA NPs with erythrocyte membranes to adopt longcirculating stealth properties in the bloodstream.78 The latter is in part mediated by the cell surface
expression of specific membrane proteins that function as ‘markers of self’, such as the CD47
glycoprotein that inhibits MPS clearance via binding of the signal regulatory protein- (SIRP-) on
phagocytes.208,209 The RBC membrane-coated polymeric NPs were prepared via a two step process in
which PLGA NPs are mechanically extruded together with preformed erythrocyte membrane-derived
nanovesicles to induce fusion of the latter on the NP surface (Figure 16). Following tail vein injection
of this novel formulation, it was demonstrated that the circulation half-life of the RBC membranecoated PLGA NPs was prolonged ~2.5 fold over PLGA NPs functionalized with conventional PEGylated
lipids.78 PEG-conjugated phospholipids, in itself also mimics of a cell’s glycocalix, are still regarded as
48
the gold standard for the engineering of lipid-based stealth nanomedicines. However, it has indeed
become clear that a hydrophilic polymer brush coating cannot completely prevent the adsorption of
serum proteins capable of inducing phagocytosis and accelerated blood clearance.208,210-212
Moreover, activation of complement has been described for PEG-phospholipid conjugates,213,214
indicating that a hydrophilic PEG shell does not guarantee immunological inertness and warranting
further investigation of better and safer alternatives for NP surface engineering. Exploiting the longcirculating properties of the RBC bio-mimetic nanocomposites, the authors very recently could show
their biomedical value in blood detoxification through specific absorption of membrane damaging
pore-forming toxins into the erythrocyte shell.215
Figure 16. Shielding of PLGA nanoparticles with erythrocyte-derived lipid membranes. Repeated
extrusion steps on a physical mixture of preformed nanoparticles and empty red blood cell (RBC)derived membrane vesicles results in RBC lipid-bilayer coating of the PLGA nanoparticle cores.
(Reprinted with permission from ref 78. Copyright 2011).
The decoration of drug loaded NPs with a bio-inspired lipid bilayer may also be advantageous toward
drug delivery on the intracellular level. De Backer et al. recently reported on the coating of siRNA
loaded biodegradable polysaccharide nanogels with natural-derived pulmonary surfactant.216,
49
Pulmonary surfactant, covering the entire alveolar surface of mammalian lungs, may potentially
interact with nanomedicines upon inhalation therapy and deep lung deposition.217 It is therefore of
key importance to carefully evaluate the effect that alveolar surfactant might have on the biological
performance of drug-loaded NPs. To this end, the authors first constructed stable surfactant-covered
siRNA nanogels. Although the coating with pulmonary surfactant substantially inhibited cellular
internalization of the nanogels in lung epithelial cells and alveolar macrophages, their gene silencing
potential in both cell types was maintained or even improved. These intriguing data suggest that
pulmonary surfactant may enhance the fraction of internalized siRNA that is delivered into the
cytosol, possibly mediated via a specific endocytic entry route or distinct intracellular trafficking
following cellular uptake.216 Accordingly, the RNAi and gene transfer activity of nucleic acid loaded
biodegradable NPs, fabricated from diamine modified poly(vinyl alcohol) grafted with PLGA, was
significantly improved in the presence of pulmonary surfactant.218,219 Here, the ternary
nanocomposites were formulated by a single-step solvent displacement technique, using the lung
surfactant as surface altering component during the preparation procedure. The latter reports mainly
ascribe the improved siRNA/DNA delivery to the surfactant-enhanced cellular internalization, as
opposed to the observations by De Backer et al..
Biodegradable antigen-loaded PLGA micro-and nanoparticles have been widely investigated in
vaccination regimens, mimicking the particulate nature of invading microorganisms to stimulate
internalization by antigen presenting cells (APCs).69,220 Moreover, the controlled release of the
encapsulated antigens enables tight control over the duration of antigen exposure in order to
minimize tolerance and stimulate an effective and long-lasting immune response.69,221,222 Conversely,
coupling the antigens to the surface of preformed particles shows better resemblance with the
natural multivalent antigen presentation on the surface of microbial pathogens. Building further on
this biomimetic antigen delivery concept, Bershteyn et al. designed phospholipid enveloped PLGA
micro-and nanoparticles with the model antigen covalently linked to PEGylated lipids stably anchored
in the phospholipid shell.52 In an attempt to enhance the immune response, the lipophilic danger
50
signals monophosphoryl lipid A (MPLA) and -galactosylceramide (-GC) were additionally inserted
in the lipid coating as adjuvants.223,224 Both lipid-coated PLGA micro-and nanoparticles evoked high
titers of antigen-specific IgG using only nanogram antigen doses, which could be even further
improved by the pathogen-mimetic presentation of the above-mentioned lipophilic adjuvants.
Moreover, the reported particulate vaccination platform significantly outperformed a conventional
vaccination strategy with soluble protein antigen combined with the particulate adjuvant alum or a
mixture of MPLA/-GC as molecular adjuvants.52 Similarly, as briefly mentioned in section 4.3.,
SMBVs were also proposed as particulate antigens. Herein the charged polysaccharide core could
serve for encapsulation of (intracellular) antigens while membrane-associated antigens could be
integrated in the surrounding lipid bilayer, thus imitating the natural presentation of antigenic
epitopes on the surface of microbial pathogens. In addition, the incorporation of nucleic acidderived, protein-derived and lipophilic adjuvants has been suggested of which the latter is also
assumed to be inserted in the lipid layer of the SMBV.54
5. TOXICOLOGICAL CONSIDERATIONS OF LIPID-POLYMER MATRIX NANOPARTICLES
It is becoming increasingly clear that careful NP design is key to reach the goal of both safe and
effective nanomedicinal drug delivery. Although the major goal of drug encapsulation in targeted
nanomedicines is to reduce off-target toxicity, the toxicity inferred by the nanocarrier itself cannot
be ignored and also requires thorough assessment. The NP composition, physicochemical properties
(size, shape, porosity and surface charge) as well as residual chemical or biological impurities may
contribute to a toxic response. Nanotoxicity can be evoked via many distinct mechanisms including,
but not limited to, the induction of reactive oxygen species (ROS), cellular membrane damage and
immunotoxicity. Below, several important findings are highlighted in the context of lipid-polymer
hybrid NPs. For a more detailed insight in the field of nanotoxicology the reader is referred to several
excellent and comprehensive reviews available in the literature.225-229
51
It is known that upon in vivo administration, NPs will interact with and bind various proteins, thereby
creating a so-called protein corona on their surface.182 The identity of this corona is mainly influenced
by the nature of the microenvironment (blood circulation, interstitial fluid, cytosol,…) and the NP
physicochemical characteristics.230 Hydrophobic and charged NPs are especially prone to the
deposition of a stable protein corona upon immersion in biological fluids.227,230 Importantly, the
adsorption of proteins to a NP surface will influence its interaction with target cells and modulate NP
internalization.182 The cellular uptake mechanism of NPs might be related to the observed toxic
response since it determines the intracellular localization of the NP as well as the total (intra)cellular
dose. As a consequence, factors that affect the extent and/or the mechanism of cellular
internalization, e.g. the presence of a protein corona, likely also influence the level of nanotoxicity.225
Additionally, the proteins that interact with NPs might undergo conformational changes, thereby
impairing the biological function or inducing NP recognition by cells of the immune system.40,225
Likewise, the induction of NP aggregation through interparticle protein bridging as well as NP
opsonization contribute to NP scavenging by MPS phagocytes in vivo, possibly leading to
immunotoxicity.40 It is conceivable that endowing the polymeric NPs with a neutral lipid coat will
improve the overall tolerability of the formulation, as it obscures the hydrophobicity and charge of
the underlying NP. Adding PEG-lipids to the lipid-envelope is believed to further enhance the
biocompatibility by mitigating unspecific interactions and decreasing phagocytic clearance.40
However, immunotoxicity still remains an issue given that all liposomes have the intrinsic potential of
inducing complement activation, possibly leading to type I hypersensitivity reactions and fast MPS
clearance, with negatively or positively charged lipid bilayers showing the highest reactivity.231
Further underscoring the importance of charge, Kedmi et al. documented significant toxicity in vivo
in mice following intravenous injection of cationic lipids, as demonstrated by a marked hepatotoxic
response and weight loss, in contrast to their neutral and negatively charged counterparts.232 As
briefly mentioned in section 4.4., even the insertion of PEGylated lipids can trigger complement
activation.231,233 Moghimi et al. demonstrated that the negatively charged phosphodiester group
52
within the PEG-phospholipid was responsible for this effect, since complement activation could be
prevented by methylation of this anionic moiety or by application of neutral PEG-lipid
conjugates.234,235 Furthermore, it has been demonstrated that the intravenous injection of a first dose
of PEGylated liposomes could trigger the emergence of anti-PEG IgM, which initiate rapid clearance
of a second liposome dose.212,236,237 This so-called accelerated blood clearance (ABC) phenomenon
has recently been reviewed in detail.238 In addition, decorating the surface of (PEGylated) lipidpolymer nanohybrids with targeting ligands and/or therapeutics might further add to the concern of
immune reactivity.181
Also at the cellular level, the size and the surface characteristics of NPs are the predominant factors
that dictate interactions with cells or cellular components and hence drive toxicity.226 Many drug
loaded nanoparticles carry a net positive surface charge to allow electrostatic interaction with the
negatively charged cell membrane and to trigger adsorptive endocytosis. However, it has been
demonstrated that the attachment of positively charged matrix NPs to the plasma membrane can
lead to its concentration dependent deformation and rupture. The resulting influx of extracellular
calcium may aggravate cellular toxicity by contributing to mitochondrial dysfunction and
proapoptotic signaling.239 Induction of membrane damage at the level of erythrocytes can cause
significant hemolysis, e.g. as demonstrated for silica NPs as a function of the surface density of
available silanol (SiOH) groups, NP porosity and surface charge.240,241 MSNPs are believed to be less
toxic than their solid non-porous counterparts, given that the controlled porosity reduces the contact
area with biological membranes and lowers the surface density of silanol groups, which are believed
to be largely responsible for the observed toxicity.105 Given the involvement of the surface
characteristics of the NPs in the observed cellular toxicity, one can anticipate that masking the
surface of the polymeric NPs by coating them with neutral or oppositely charged phospholipids could
alleviate their cytotoxicity. Again, the use of cationic lipids for coating should be avoided given that
transfer of cationic lipids to the plasma membrane can equally cause membrane destabilization, Ca2+
influx and oxidative stress.116 Moreover, it has been demonstrated that cationic lipids can evoke a
53
pro-inflammatory response through activation of toll-like receptor 4 (TLR4).232 As highlighted earlier,
many drugs require cytosolic delivery (e.g. nucleic acid therapeutics) and therefore depend on the
efficiency with which their nanocarrier is able to mediate escape from the endosomal compartments
following endocytic internalization. Endosomal escape is typically accompanied by rupture or
permeabilization of the endosomal/lysosomal membrane, thereby releasing the lysosomal protease
cathepsin B which is known to trigger inflammasome activation and apoptosis.242-244 The impact of
cationic lipid-based nanomedicines on intracellular signaling pathways and cellular toxicity has
recently been summarized in a comprehensive review by Lonez et al.245 Cytosolic accumulation of
polycationic material can in turn interfere with protein function or induce impairment of cellular
organelles.246
To avoid accumulation of polymeric matrix NPs in cells and tissues upon repeated administration, it is
of equal importance that they are composed of biodegradable constituents. For instance, as
mentioned in section 3.1., PLGA and PLA NPs can be fully degraded via ester hydrolysis in the
toxicologically acceptible products lactic and glycolic acid.71 MSNPs show faster dissolution kinetics
compared to solid non-porous Stöber silica NPs owing to the higher total surface area, leading to
nontoxic soluble silicic acid products.105 Hydrogel NPs can be rendered biodegradable by selecting
(enzymatically or hydrolytically) degradable polymer building blocks or incorporating degradable
crosslinks into the 3D polymeric network.120 Of note, the possible adverse biological interactions
mediated by the emerging (polymeric) degradation products should evenly be taken into account.
As nanomedicines will interact with various cells and biomolecules upon in vivo administration, it is
crucial that the toxicological impact of these interactions is carefully assessed. Summarizing the
findings above, the potential toxicity of lipid-coated polymer NPs likely correlates well with what has
been independently described for their lipid and polymer constituents. As a general rule, the use of
charged (mainly polycationic) components is preferably avoided. Although the use of a (PEGylated)
lipid-coat can offer additional protection at the extracellular level and the level of the cell membrane,
activation of complement and resulting immunotoxicity cannot be completely ruled out. It is
54
therefore advisable to meticulously document the nanotoxicological profile for every new NP design
prior to clinical application, in order to identify potential side effects. To this end, novel standardized
and high-throughput methodologies that allow a multiparametric assessment of (cellular) toxicity
and in vitro models that enable a more reliably prediction of in vivo toxicity could prove highly
valuable.225,228,229,239
6. CONCLUSIONS AND FUTURE OUTLOOK
A plethora of nanoparticles (NPs) with diverse nanoarchitectures has been witnessed to date with
the aim to improve the efficacy and safety of encapsulated drugs. Biocompatible nanomedicines that
enable a more selective delivery of therapeutics to diseased cells while mitigating exposure of
healthy tissues, are highly sought after. However, overcoming the countless barriers a drug-loaded
NP encounters en route to its biological target remains a daunting task. Both polymeric NPs and
liposomes have already demonstrated in vivo potential, but still suffer from inherent shortcomings.
Capitalizing on recent advances in nanotechnology together with a more fundamental insight in the
behavior of NPs upon in vivo administration, should pave the way to the engineering of better and
safer nanomedicines. Lipid-enveloped polymer matrix NPs mainly aim to merge the advantages of
both materials in a single nanocarrier and at the same time tackle their limitations. These lipidpolymer nanohybrids provide a flexible platform affording ample control over their physical,
chemical and biological attributes. This degree of flexibility is of the utmost importance to bypass the
numerous extra-and intracellular biological hurdles and to provide a suitable drug delivery solution
for the ever expanding collection of pharmaceuticals and adjuvants with greatly diverging
physicochemical properties. In addition, the complex pathophysiology encountered in various
diseases also demands a more complex integrative drug delivery approach. Lipid-polymer hybrid NPs
have proven to be particularly suited toward targeted drug combination therapy in order to enhance
therapeutic efficacy and reduce drug resistance. As polymeric NPs and liposomes have already been
widely investigated for stimuli-responsive and image-guided drug delivery, also these traits can be
55
implemented in the engineering of lipid-polymer hybrids to gain spatiotemporal control over the
drug delivery process and further expand their use in theranostic applications.
On the other hand, the currently witnessed transition toward more sophisticated multifunctional
nanocarriers also entails novel challenges that may hamper clinical translation. From a
manufacturing perspective, questions arise on the feasibility to scale-up the often labor-intensive
multistep production process while meeting quality control standards and safeguarding costeffectiveness. Most nanoformulations are designed for parenteral administration and thus require
aseptic production, which could further add to the overall production costs.19 Therefore, it remains
imperative to devote ample attention to clinical benefit by evaluating how these innovative drug
delivery systems could outperform current state-of-the-art nanomedicines. An important step
forward is the development of single-step and continuous (microfluidic) synthesis of lipid-coated
polyester NPs via nanoprecipitation, thereby reducing batch-to-batch variation while maintaining a
sufficiently high production rate. However, this might not be as feasible for other types of lipidcoated nanomaterials and likely requires independent optimization. To ensure large-scale
pharmaceutical production, the horizontal scaling-out of the production process through connection
of multiple small-scale production units running in parallel could be favored over conventional scaleup of a single production unit.14
In the advent of nanotoxicology, it will also become increasingly important to make an in-depth
assessment of the pharmacokinetic, pharmacodynamic and (immuno)toxicological profile. It is
conceivable that multicomponent nanocarriers, constructed from materials with divergent
properties, might pose higher risk for carrier-mediated off target toxicity. The emphasis should
therefore be put as much as possible on biocompatible and biodegradable materials. The latter may
also spur future investigation of bio-mimetic lipid-polymer nanocarriers as a biofunctional yet
biocompatible alternative.
Altogether, as judged by the emerging research output on lipid-polymer nanocomposites, these
hybrid nanocarriers evoked much enthusiasm among the drug delivery community. With further
56
research investments on the level of manufacturing and pre-clinical assessment in appropriate
animal models, we are confident that these multifaceted drug delivery systems could have broad
clinical impact.
ACKNOWLEDGEMENTS
KR is a postdoctoral fellow of the Research Foundation-Flanders (FWO-Vlaanderen). Financial
support by the Ghent University Special Research Fund and the FWO is acknowledged with gratitude.
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