Modular construction of multifunctional bioresponsive cell-targeted nanoparticles for gene delivery

Article abstract of DOI:10.1021/bc100149g

Multifunctional and modular block copolymers prepared from biocompatible monomers and linked by a bioreducible disulfide linkage have been prepared using a combination of ring-opening and atom-transfer radical polymerizations (ATRP). The presence of termi

Full text of DOI:10.1021/bc100149g

ARTICLE
pubs.acs.org/bc
Modular Construction of Multifunctional Bioresponsive Cell-Targeted
Nanoparticles for Gene Delivery
Aram O. Saeed, Johannes P. Magnusson, Emilia Moradi, Mahmoud Soliman,§ Wenxin Wang,‡ Snow Stolnik,
Kristofer J. Thurecht,|| Steven M. Howdle,† and Cameron Alexander*
†School of Pharmacy and School of Chemistry, University of Nottingham, University Park, Nottingham NG7 2RD, United Kingdom
‡Network of Excellence for Functional Biomaterials, National University of Ireland, Galway, Ireland
§Department of Pharmaceutics, Faculty of Pharmacy, Ain Shams University, Monazamet El Wehda El Afrikia Street, El Abbassia,
Cairo, Egypt
Australian Institute for Bioengineering and Nanotechnology (AIBN), University of Queensland, Brisbane, Queensland 4072, Australia
S Supporting Information
b
ABSTRACT: Multifunctional and modular
block copolymers prepared from biocompati-
ble monomers and linked by a bioreducible
disulfide linkage have been prepared using a
combination of ring-opening and atom-
transfer radical polymerizations (ATRP).
The presence of terminal functionality via
ATRP allowed cell-targeting folic acid
groups to be attached in a controllable
manner, while the block copolymer architec-
ture enabled well-defined nanoparticles to be
prepared by a water-oil-water double
emulsion procedure to encapsulate DNA with high efficiency. Gene delivery assays in a Calu-3 cell line indicated specific folate-
receptor-mediated uptake of the nanoparticles, and triggered release of the DNA payload via cleavage of the disulfide link resulted in
enhanced transgene expression compared to nonbioreducible analogues. These materials offer a promising and generic means to
deliver a wide variety of therapeutic payloads to cells in a selective and tunable way.
’ INTRODUCTION
and accordingly, several groups have prepared block copolymers
or conjugates wherein the PEG block is connected to the hydro-
phobic component or drug via cleavable or reversible/disrupta-
ble linkers.17-19 This enables the PEG-corona to separate from
the core, exposing the contents for release, in a manner analogous
to the shedding of viral coat proteins. PEGylated or other
copolymer micelles have been successfully decorated with cell
targeting functionality,20 for example, by attaching ligands such
as folic acid derivatives to the outside of the PEG layer. The
overexpression of the folic acid receptor on certain cancer cells
has been explored by a number of groups, who have end-
functionalized polymer drug delivery systems with folic acid or
folates to enhance uptake via folate-receptor mediated endocy-
tosis.13,21,22
Carrier materials for therapeutic biopolymers must fulfill a
number of demanding criteria,1 and for nucleic acid delivery, the
hurdles are especially challenging.2-4 DNA and RNA delivery
vehicles are required to package the polynucleotide and protect it
during transport in and across numerous biological environ-
ments and barriers, and then must release the biopolymer in the
correct cellular location. The delivery materials should also be
nontoxic, and ideally should be biodegradable. The challenges of
gene delivery have led to many sophisticated carrier systems
being developed, including block copolymer micelles, vesicles,
and nanoparticles,5-8 as the core-shell architectures enable
packaging of DNA or RNA into a protected interior phase.
Further derivatization has allowed the generation of finely
structured nanomaterials that can display bioactive ligands for
cell or tissue-specific targeting above a hydrophilic corona for
enhanced biodistribution.9 Poly(ethylene glycol) (PEG) block
copolymer micelles have been a particular focus for drug and
gene delivery,10-16 utilizing the ability of PEG to reduce
opsonization and clearance from the body. Increased stability
via PEGylation can have the downside of reducing cell uptake,
However, the task of assembling these complex polymeric and
supramolecular assemblies has hitherto required equally complex,
often multistep, synthetic strategies. We describe here how the
Received: March 24, 2010
Revised:
December 1, 2010
Published: January 11, 2011
r
2011 American Chemical Society
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Scheme 1. Schematic Illustration of (a) Preparation of Folate Functionalized Nanoparticles; (b) Accumulation of Folate
Functionalized Nanoparticles in Interstitial Space via the EPR Effect; (c) Folate Receptor Mediated Endocytosis, Shell
Detachment, and Core Erosion of the Particles to Enhance Intracellular Release of Plasmid DNA
judicious design of functional initiators combined with ring-
opening polymerization and the activators generated by electron
transfer atom transfer radical polymerization (AGET ATRP)
methodologies can be adapted to assemble modular, multicom-
ponent polymers in an experimentally facile manner. Control of
the overall structure is established at the start through the use of a
reductively cleavable initiator to grow two different polymer
blocks, with an easily accessible chain-end for ligand fictionaliza-
tion. The key to this strategy is that the chemistries of the block
copolymers so-produced enable formulation with hydrophilic
biopolymers in a water/oil/water (w/o/w) system, allowing the
incorporation of a therapeutic payload in high efficiency and,
importantly, generating correctly positioned surface and core
functionality (Scheme 1).
neutral alumina. Copper(II) bromide (CuBr2, 99%), 2-bromo-2-
methylpropionyl bromide, L-ascorbic acid (99%), N,N,N0,N00,
N000-Pentamethyldiethylenetriamine (PMDTA, 98%), and tin-
(II)-2-ethylhexanoate were used as received from Sigma Aldrich.
N,N-Azobis(isobutyronitrile) (AIBN, 98%, Aldrich) was recrys-
tallized from ethanol. Plasmid DNA GWIZLuc, containing the
firefly luciferase gene, was supplied as a 5 mg/mL solution by
Aldevron (South Fargo, USA) and used without further treat-
ment. The luciferase detection kit was acquired from Promega
(Southampton,UK) and contained cell culture lysis reagent
(5ꢀ solution), Luciferin reagent, and Luciferase assay buffer.
Phosphate buffer saline (PBS) was used as received from Fisher
Scientific. Dialysis membrane (MWCO 3500, regenerated
cellulose) was used as received from Spectrapor.
We have exemplified the synthetic strategy with gene delivery
experiments in Calu-3 cells, which are human carcinoma cells
that produce mucus, and thus have been used as a clinically
relevant model for nasal and diseased airway epithelia.23,24
Synthesis of 2-Bromo-2-methyl-propionic Acid 2-(2-
Hydroxy-ethyldisulfanyl)-ethyl Ester (1). The heterobifunc-
tional ATRP/ROP initiator was prepared as follows. Dith-
iodiethanol (12 g, 0.077 mol) was added to a 100 mL round-
bottom flask equipped with three-way stopcocks connected to
either an argon line or a vacuum pump, and dried via azeotropic
distillation with anhydrous toluene (3 ꢀ 20 mL) under reduced
pressure. After complete evaporation of toluene, anhydrous THF
(20 mL) was added via a syringe under an argon atmosphere.
Triethylamine (2.3 g, 0.023 mol) was added to the solution
which was maintained at 4 °C during slow addition of R-
bromoisobutyryl bromide (2.6 g, 0.011 mol). The solution was
allowed to warm to room temperature and left overnight. The
solution was then filtered to remove the salt, and the crude
product was recovered by evaporation of THF. The product was
’ EXPERIMENTAL PROCEDURES
Materials. All solvents and reagents were of analytical or
HPLC grade and purchased from Sigma or Fisher Scientific
unless otherwise stated. Deuterated solvents were from Sigma.
The two monomers DL-lactide and glycolide were obtained from
Purac biochem. The monomers were recrystallized from ethyl
acetate prior to use. Poly(ethylene glycol) methyl ether
(PEGMA-ME 475, Mn 475) were purchased from Sigma Aldrich
and purified before use by passing through a column filled with
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ARTICLE
purified by passing through silica column chromatography using
ether as eluent.
NMR (CDCl3): δ = 169.5, 166.3, 71.9, 70.5, 59.2, 16.9. FTIR
(cm-1): 2860-2940 cm-1 (C-H), 1760 cm-1 (CdO).
Synthesis of Azide-Terminated PLGA-S-S-PEGMA475
Block Copolymer (P2-b). Sodium azide (0.04 g, 0. 62 mmol)
was added to a round-bottom flask containing PLGA-S-S-PEG-
MA475 block copolymer (4 g, 0.3 mmol) dissolved in DMF
(12 mL). The reaction mixture was stirred at 50 °C for 24 h after
which time it was allowed to cool down to room temperature,
After elimination of DMF under reduced pressure, toluene
(15 mL) was added, and the insoluble salt was removed by
centrifugation (10 000 rpm at 25 °C for 25 min). Finally, the
toluene was evaporated under reduced pressure. The product
was characterized by FTIR, 1H NMR, and GPC.
1H NMR (CDCl3, 400 MHz)): 1H NMR (CDCl3, 400
MHz)): δ=4.48-4.44 (t, 2H, CH2-OH), 4.46 (t, J=6.68 Hz,
2H), 3.92-3.8 (m, 2H, CH2-O-CdO), 3.91 (qapparent, Japp
=
5.94, 2H), 3-2.97 (t, 2H, CH2-S), 2.98 (t, J = 6.62 Hz, 2H),
2.91-2.89 (t, 2H, CH2-S), 2.90 (t, J =5.835, 2H), 1.95 (s, 6H,
CH3), 2.1 (t, J = 6.185, H). 13C NMR (CDCl3): δ=171.9, 63.7,
60.4, 55.6, 41.5, 36.6, 30.8. HR-MS (ESþ) calcd for
C8H15BrO3S2 (MHþ) 301.965 and 303.237, found 301.8102
and 303.8056. FTIR (ν, cm-1): 3410 (-OH), 2927 (C-H),
1734 (CdO).
Synthesis of PLGA Polymer Using 2-Bromo-2-methyl-
propionic Acid 2-(2-Hydroxy-ethyldisulfanyl)-ethyl ester
(P1). PLGA, Poly(DL-lactide-co-glycolide) was synthesized by
ROP (ring-opening polymerization) using 2-((2-hydroxyethyl)
disulfanyl)ethyl 2-bromopropanoate as initiator. First, DL-lactide
(1.82 g, 0.012 mol), glycolide (1.47 g, 0.012 mol), and 2-((2-
hydroxyethyl)disulfanyl)ethyl 2-bromopropanoate (0.4 g, 0.0-
014 mol) were added to a polymerization tube and purged with
argon several times before being heated in an oil bath to 80 °C
and purged with argon for a further 2 h. The reaction mixture was
then heated to 140 °C in order to conduct the polymerization in
the melt phase. At this point, tin(II) 2-ethylhexanoate (0.106 g,
0.27 mmol) was added, and the mixture was stirred for 24 h. The
polymerization tube was then removed from the oil bath and
cooled down to room temperature. The polymer was recovered
by dissolution in THF and precipitated in methanol three times
to remove unreacted monomer. The precipitate was filtered and
dried under reduced pressure to give a white product (yield
95%). The molecular weight of the polymer and polydispersity
index were determined by gel permeation chromatography using
chloroform as eluent.
1H NMR (CDCl3, 400 MHz). δ = 5.18 (m, 1H, CH-CdO),
4.82 (m, 2H, CH2-CdO), 3.94-4.14 (m, CH2, CH2-O-CdO),
3.8-3.4 (m, CH2CH2O), 3.2-3.38 (s, CH3), 1.58 (s, 3H, CH3),
1.6-1.7 (m, CH2-N3), 0.7-1.3 (s, CH3). FTIR (ν, cm-1):
2860-2940 (C-H), 2106 (N3), 1760 (CdO).
Synthesis of Propargyl Folate. This was accomplished by a
method modified from the literature.22 Folic acid (1.0 g, 2.2
mmol) was dissolved in DMF (10 mL) and cooled in a water/ice
bath. N-Hydroxysuccinimide (0.26 g, 2.5 mmol) and 1-ethyl-
3-[3-dimethylaminopropyl]carbodiimide hydrochloride (EDC,
0.44 g, 2.5 mmol) were added, and the resulting mixture was
stirred in an ice bath for 30 min to give a white precipitate.
A solution of propargylamine (0.124 g, 2.25 mmol) in DMF
(5.0 mL) was added, and the resulting mixture was allowed to
warm to room temperature and stirred for 24 h. The reaction
mixture was poured into water (100 mL) and stirred for 30 min
to form a precipitate. The orange-yellow precipitate was
filtered, washed with acetone, and dried under vacuum to give
a product in 90% yield. Analytical data corresponded to those of
the literature reported product.22
1H NMR (DMSO-d6, 400 MHz): 8.64 (s, 1H, PtC7 H),
8.29-8.24 (d, 1H, PtC6-CH2NH-Ph), 8.04-8.02 (d, 1H, -
CONHCHCO2H), 7.67-7.65 (d, 2H, Ph-C2H and Ph-C6H),
6.93 (br s, 2H, NH2, 2H), 6.65-6.63 (d, 2H, Ph-C3H and Ph-
C5H), 4.49-4.48 (d, 2H, PtC6-CH2NH-Ph), 4.32-4.30 (m,
1H, -CONHCHCO2H), 3.84-3.81 (m, 2H, -CONH-
CH2CtCH), 3.07-3.05 (t, 1H, -CONHCH2CtCH), 2.88
(s, 1H, -CONH-CH2CtCH,), 2.72 (br s, 1H, OH,), 2.31-
2.20 (m, 2H, -CH2CO2H), 1.98-1.96 (m, 1H, -CHCH2-
CH2), 1.87-1.85 (m, 1H, -CHCH2CH2). Note: Pt = pteridine.
“Click” Reaction of Azide Terminal of PLGA-S-S-pPEG-
MA475 Blocks Copolymer with Propargyl Folate (P3-a). In
a 25 mL round-bottom flask, PLGA-S-S-PEGMA475-N3 (2 g,
1H NMR (CDCl3, 400 MHz)): δ = 5.18 (m, 1H, CH-CdO),
4.82 (m, 2H, CH2-CdO), 4.2-4.3 (m, 4H, CH2-O-CdO),
2.8-2.9 (m, 4H, CH2-S), 1.8-1.9 (s, 6H, CH3), 1.58 (d, 3H,
CH3-CH). 13C NMR (CDCl3): δ = 169.5, 166.3, 69.1, 60.8,
53.5, 16.9. FTIR (cm-1): 2860-2940 cm-1 (C-H), 1760 cm-1
(CdO).
Synthesis of PLGA-S-S-pPEGMA475 Block Copolymer
Using (P2-a). The PLGA-S-S-PEGMA475 block copolymers
were prepared by atom transfer radical polymerization (ATRP)
using the PLGA-S-S-Br block as macroinitiator. For PLGA-S-S-
PEGMA475 block copolymers, PLGA-S-S-Br (1.5 g, 0.3 mmol,
poly(ethylene glycol)methyl ether methacrylate 475 (2.8 g, 6.3
mmol), CuBr2 (0.067 g, 0.3 mmol), and PMDTA (63 μL, 0.3
mmol) were added into a round-bottom flask and the flask
purged with argon several times, followed by addition of buta-
none (10 mL) and degassing for 20 min under an argon atmo-
sphere. L-Ascorbic acid (0.022 g, 0.12 mmol) was added to the
polymerization solution under argon. The mixture was then
heated in an oil bath to 50 °C and stirred for 3.5 h. The poly-
merization product was precipitated in a mixture of hexane/ether
(50:50), and then the polymer was passed through a short
alumina column to remove copper salts. Finally, the solvent
was evaporated and the product dried under reduced pressure.
The product was characterized by NMR and GPC using chloro-
form as eluent.
0.2 mmol, Mn = 9000 g mol-1), PMDTA (0.034 g, 0.02
3
mmol), and propargyl folate (0.014 g, 0.24 mmol) were dis-
solved in DMF (10 mL) and purged with argon for 30 min. The
solution was then transferred via a degassed cannula needle to
another tube equipped with a magnetic stir bar containing
CuBr (0.028 g, 0.2 mmol) under argon atmosphere. The
mixture was stirred at 26 °C for 24 h in the absence of oxygen.
After the reaction, the mixture was exposed to air, and the
solution was passed through a column of neutral alumina. DMF
was removed under vacuum. The resulting folate-terminated
PLGA-S-S-pPEGMA475 was dissolved in THF and filtered to
remove the excess of propargyl folate. The product was further
purified by dialysis in acetone using regenerated cellulose
membrane molecular weight cutoff of 3500 Da. Finally, the
polymer was dried under vacuum for 6 h, and the product was
1H NMR (CDCl3, 400 MHz). δ = 5.18 (m, 1H, CH-CdO),
4.82 (m, 2H, CH2-CdO), 3.94-4.14 (m, CH2,
CH2-O-CdO), 3.4-3.8 (m, CH2CH2O), 3.2-3.38 (s, -CH3),
1.58 (s, 3H, CH3), 1.6-1.9 (m, CH2), 0.7-1.3 (s, CH3). 13C
1
characterized by FTIR, H NMR, and GPC. The synthesis of
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P3-b was exactly analogous except propargyl alcohol was used in
place of propargyl folate.
CH2CH2-O-CdO), 1.58 (d, 3H, CH3-CH). 13C NMR
(CDCl3): δ = 169.5, 166.3, 69.1, 60.8, 53.5, 16.9. FTIR (cm-1
)
P3-a: 1H NMR (CDCl3, 400 MHz) δ=8.5, 8.2, 7.5, 6.5 (weak
multiplets, folate terminus), 5.18 (m, 1H, CH-CdO), 4.82 (m,
2H, CH2-CdO), 3.94-4.14 (m, CH2, CH2-O-CdO),
3.8-3.4 (m, CH2CH2O), 3.2-3.38 (s, -CH3), 1.58 (s, 3H,
CH3), 1.6-1.9 (m, CH2), 0.7-1.3 (s, CH3)). 13C NMR
(CDCl3): δ = 169.5, 166.3, 71.9, 70.5, 59.2, 16.9. FTIR
(cm-1) 2860-2940 cm-1 (C-H), 1760 cm-1 (CdO).
P3-b: 1H NMR (CDCl3, 400 MHz) δ = 5.18 (m, 1H,
CH-CdO), 4.82 (m, 2H, CH2-CdO), 3.94-4.14 (m, CH2,
CH2-O-CdO), 3.8-3.4 (m, CH2CH2O), 3.2-3.38
(s, -CH3), 1.58 (s, 3H, CH3), 1.6-1.9 (m, CH2), 0.7-1.3 (s,
CH3)). 13C NMR (CDCl3): δ = 169.5, 166.3, 71.9, 70.5, 59.2,
16.9. FTIR (cm-1) 2860-2940 cm-1 (C-H), 1760 cm-1
(CdO).
2860-2940 cm-1 (C-H), 1760 cm-1 (CdO).
Fabrication of Nanoparticles. Nanoparticles with PLGA-S-
S-PLGA (P4):PLGA-S-S-PEGMA475(P3) ratio of 50:50 were
prepared using a modified solvent diffusion technique. P3 and P5
(50 mg of each) were dissolved in dichloromethane (2 mL). The
organic solution was then mixed by vortex agitation for 30 s
with an aqueous phase (200 μL) of DNase-free water with or
without plasmid DNA (200 μg). The obtained emulsion was
then poured into a polar phase (25 mL of ethanol) under
moderate magnetic stirring, resulting in an immediate precipita-
tion of polymer nanoparticles. The suspension was diluted with
ultrapure water (25 mL) and stirred for 1 h. The organic solvent
was evaporated under reduced pressure at 30 °C. The final
particles suspension was passed through a Sephadex column (PD
10) and then characterized by dynamic light scattering (DLS),
zeta potential measurements, and transmission electron micro-
scopy (TEM).
Synthesis of PLGA-S-S-PLGA Homopolymer (P4) Using
2,2-Dithiodiethanol as Bifunctional Initiator. PLGA-S-S-
PLGA homopolymer, poly(DL-lactide-co-glycolide-s-s-DL-lac-
tide-co-glycolide) was synthesized by ROP (ring-opening poly-
merization) using dithiodiethanol (2) as bifunctional initiator.
First, DL-lactide (2.7 g, 0.018 mol), glycolide (2.1 g, 0.018 mol),
and dithiodiethanol (0.3 g, 0.0019 mol) were added to a
polymerization tube and purged with argon several times before
being heated in an oil bath to 80 °C and purging with argon for a
further 2 h. The reaction mixture was then heated to 140 °C in
order to conduct the polymerization in the melt phase. At this
point, tin(II) hexanoate (0.078 g, 0.2 mmol) was added, and
the mixture was stirred for 24 h. The polymerization tube was
then removed from the oil bath and cooled down to room
temperature.
’ CHARACTERIZATION OF NANOPARTICLES
Dynamic Light Scattering. The size and polydispersity of
nanoparticle suspensions were determined by dynamic light
scattering. An aliquot (70 μL in water) was added in a quartz
cuvette and the sample was examined on a Viscotek 802 dynamic
light scattering instrument. Scattered light from a 50 mW laser
source (830 nm) was recorded from an internal light detector
aligned at 90° from source. From standard autocorrelation
functions, measured correlation coefficients were related to
hydrodynamic radii at varying temperatures via the Stokes-
Einstein equation, RH = KT/6πηD, where RH is the hydrody-
namic radius, k is the Boltzmann constant, T is the temperature,
and η is the viscosity of the solvent. The size distribution of the
scattering particles was determined by the average mass number
of the particles.
The polymer was recovered by dissolution in THF and
precipitated in methanol (this was repeated three times to
remove unreacted monomer). The precipitate was filtered and
dried under reduced pressure to leave a white solid product (95%
yield).
1H NMR (CDCl3, 400 MHz)): δ=5.18 (q, 1H, CH-CdO),
4.82 (q, 2H, CH2-CdO), 4.36-4.4 (m, 4H, CH2-O-CdO),
2.87-2.96 (m, 4H, CH2-S), 1.58 (d, 3H, CH3-CH). 13C
NMR (CDCl3): δ = 169.5, 166.3, 69.1, 60.8, 53.5, 16.9. FTIR
(cm-1) 2860-2940 cm-1 (C-H), 1760 cm-1 (CdO).
Synthesis of PLGA-TEG-PLGA Homopolymer (P5) Using
Tetra(ethylene glycol) (TEG) as Bifunctional Initiator. Poly-
(DL-lactide-co-glycolide-TEG-DL-lactide-co-glycolide) was synth-
esized by ROP (ring-opening polymerization) using tetra-
(ethylene glycol) initiator (3). First, DL-lactide (2.1 g, 0.014
mol), glycolide (1.7 g, 0.014 mol), and tetra(ethylene glycol)
(0.3 g, 0.0015 mols) were added to a polymerization tube and
purged with argon several times before being heated in an oil bath
to 80 °C and purging with argon for a further 2 h. The reaction
mixture was then heated to 140 °C in order to conduct the
polymerization in the melt phase. At this point, tin(II) 2-ethyl-
hexanoate (0.062 g, 0.15 mmol) was added, and the mixture was
stirred for 24 h. The polymerization tube was then removed from
the oil bath and cooled down to room temperature. The polymer
was recovered by dissolution in THF and precipitated in
methanol (this was repeated three times to remove unreacted
monomer). The precipitate was filtered and dried under reduced
pressure to yield a white solid product.
Surface Charge Measurement (Zeta Potential). The sur-
face charge of the particle was measured on a Malvern Instru-
ments Zetasizer 2000. To determine the zeta potential, the
measurements were performed in 1 mM phosphate buffer solu-
tion at pH 7.4. Mean values were obtained from three different
batches, each measured three times.
Microscopy. The morphology of the nanoparticles was ex-
amined by transmission electron microscopy (TEM). Samples
were imaged using a JEOL (JEM-1010) electron microscope. A
few drops were added onto the copper grid and allowed to dry in
air. The particles were negatively stained with 3% phosphotungs-
tic acid solution.
Determination of Plasmid DNA Encapsulation Efficiency.
The theoretical loading of plasmid DNA was 0.4% (w/w) in
regard to the total amount of P1 in the nanoparticles. For
determination of the actual amount of encapsulated DNA, an
extraction method was used. Briefly, nanoparticles (10 mg)
loaded with plasmid DNA were resuspended in 0.5 N NaOH
(1.5 mL) under horizontal shaking at 37 °C until a clear solution
was obtained. The DNA content in the resulting solution was
quantified by a UV spectrophotometry (Shimadzu 1204, Tokyo,
Japan) at 260 nm by means of a standard absorption curve
derived from DNA in water. The encapsulation efficiency
was calculated as the ratio between the actual and theoretical
DNA loading percentages. Each experiment was performed in
triplicate.
1H NMR (CDCl3, 400 MHz): δ = 5.18 (m, 1H, CH-CdO),
4.82 (m, 2H, CH2-CdO), 4.27-4.34 (m, 2H, CH2-O-CdO),
3.6-3.65 (s, 8H, CH2CH2-O), 3,67-3.72 (m, 4H,
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Release Profile of Plasmid DNA from the Nanoparticles.
The in vitro release of DNA in the presence/absence of glu-
tathione (as a representative intracellular reducing agent) was
investigated as follows. Nanoparticles (10 mg) were suspended
in glutathione solution (1.5 mL, 5 mM) at 37 °C under con-
tinuous agitation. At various time intervals, the supernatant was
withdrawn and fresh buffer was replenished. The amounts of
plasmid DNA in the supernatant were determined by UV
spectrophotometry at 260 nm. For control experiments, phos-
phate-buffered saline (PBS) pH 7.4 was used instead of glu-
tathione solution and nonreducible PLGA nanoparticles were
also studied in the presence of GSH and PBS. The experiments
were performed in triplicate.
Cell Metabolic Activity (Viability) Assay. MTS assay was
performed to evaluate the effect of nanoparticles on cell meta-
bolic activity. Calu-3 cells were seeded on 96-well plates at a
density of 10 000 cells per well and cultured in the EMEM
medium. Prior to the test, cell medium was removed and replaced
by sample solutions containing nanoparticles in HBSS. The
samples contained (4 mg.mL-1) concentration of nanoparticles;
0.1% (v/v) Triton-X 100 and HBSS were used as a positive and
negative control, respectively. Cells were incubated under stan-
dard conditions with the samples or the controls for a period of
two hours, after which they were removed. Cells were then
washed with PBS following which the MTS assay was performed
according to the manufacturer’s instructions.
detection was performed with a luciferase detection kit. Cell lysis
buffer (300 μL) was added to each well. A pipet tip was used to
scrape all cells from the well substrate, after which the lysate was
centrifuged at 13 500 rpm for 5 min and the supernatants were
collected. The luciferase activity of each sample was measured by
gently mixing reconstituted luciferin reagent (100 μL) with cell
lysate (20 μL) in a scintillation vial insert and inserting into a
luminometer (TD 20/20, Promega).
Determination of Transgene Expression Relative to Cel-
lular Protein Content. Cellular protein contents were assayed
by the Bradford method. Bovine serum albumin (BSA) standards
were prepared from a 1 mg/mL stock solution by diluting with 1/
10 strength lysis buffer to 0.6, 0.5, 0.4, 0.3, 0.2, and 0.1 mg.mL.
Aliquots of each standard (10 μL) were transferred into a 96-
assay plate; then, 200 μL Bradford assay reagent was added, and
the A595 was measured after 15 min. From samples prepared in
the luciferase detection protocol, 20 μL of each sample was
placed in an Eppendorf tube and diluted with 180 μL of water (to
remove interference from lysis buffer). Again, aliquots of 10 μL
were transferred to the wells of a 96-well assay plate and 200 μL
Bradford reagent was added. The mixtures were allowed to stand
for the same time as the standards, and then, the UV absorption
at 295 nm (A595) was measured. A standard curve was prepared
and a linear curve fit used to give the equation relating A595
to protein concentration, which was employed to calculate the
protein concentration in the samples. Results in light units.mL-1
obtained from previous sections were divided by the total
concentration of protein in the sample (in mg.mL-1) from this
assay to give a final transgene expression yield as reported in light
units/per mg of cellular protein.
’ CELLULAR UPTAKE EXPERIMENTS
Prior to the experiment, the culture medium was removed, and
the cell monolayers rinsed with phosphate buffer saline (PBS)
and then allowed to equilibrate for 1 h at 37 °C in Hanks
balanced salt solution (HBSS) with 15 mM 4-(2-hydroxyethyl)-
1-piperazine-ethane sulfonic acid (Sigma). Series batches of
nanoparticles were prepared, varying in the mass ratio of P3a/
P4 from 5% to 75%, and a batch of P3b/P4 of 75% mass ratio was
used as control. All the batches were labeled with Coumarin 6 to
assess the cellular uptake. Nanoparticle samples composed of
2 mg.mL-1 nanoparticles suspended in HBSS buffer were
applied, and the uptake study performed for a period of 4 h.
The samples were then removed and the cell monolayer washed
with buffer. In the next stage, the cells were digested with 0.2 N
NaOH in 0.5% Triton X-100 and centrifuged at a speed of 13 000
rpm for 30 s (Eppendorf 5417 R centrifuge-AG 22331
Hamburg). The fluorescence intensity of the supernatant was
measured by fluorescence spectrophotometer (MFX Microtiter
Plate Fluorimeter, The Microtiter Company, UK).
For experiments to assess uptake specificity of the particles via
expressed folate-receptors on the Calu-3 cell surfaces, 1 mmol.L
free folate was added to the cell suspensions prior to application
of the nanoparticles.
Transgene Expression Studies. Twelve well plates were
seeded with 1 ꢀ 105 cells per well, and the cells were incubated
at 37 °C and 5% CO2 for 24 h. After 24 h, the medium was aspirated
off from the wells and gently replaced with DNA-loaded nanopar-
ticles. A volume of nanoparticles containing 4 μg of DNA was
prepared as described above. The cells were incubated for 4 h at
37 °C and 5% CO2. Following this incubation, the nanoparticles
were removed and replaced with RMPI-1640 medium and the cells
were incubated for a further 24 h before analysis.
’ RESULTS
A key motivation for this study was the limiting availability and
cost of heterobifunctional PEGs for concurrent copolymer
micellization and ligand decoration; hence, new strategies for
controlled synthesis of block copolymers and nanoparticles with
more easily functionalizable end-groups are required.25,26 Our
concept therefore was to generate surface-functionalized multi-
component core-shell drug delivery systems, containing che-
mistries to enable (a) cell surface receptor targeting, (b)
prevention of extracellular protein adsorption and opsonization,
(c) intracellularly triggered loss of the polymer shell,27,28 and,
ultimately, (d) biodegradation. In addition, we aimed to combine
block components such that the copolymers could assemble into
well-defined nanoparticles in water/oil/water (w/o/w) emul-
sions, in order to encapsulate and protect sensitive biopolymers
such as DNA. Accordingly, we chose to prepare block copoly-
mers of poly(DL-lactic-co-glycolic acid) (PLGA - the hydro-
phobic core block) and poly(ethylene glycol) methyl ether
(PEGMA475 - the hydrophilic outer block), connected by a
bioreducible disulfide bridge.29 In addition, we required facile
derivatization of the hydrophilic block terminus to install a ligand
for cell receptor targeting. The folate ligand-receptor pairing
was thus chosen, since installation of folate groups at end-groups
of polymers prepared by atom transfer radical polymerization
(ATRP) has been carried out with high control and good
yields.30 ATRP offers an additional advantage in that the many
commercially available acrylic ester monomers can be polymer-
ized with a high degree of control.31,32 A number of poly-
(ethylene glycol) methacrylates (PEGMAs) can be obtained
that vary in side-chain length and water solubility, enabling
Detection of Luciferase. The medium was removed from
the cells, which were then washed with PBS twice. Luciferase
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Scheme 2. Synthetic Strategy for End-Functional Block Copolymers
control not only of molar mass and polydispersity via ATRP, but
also the hydrophilicity and architecture of the surface corona of a
micelle or vesicle.33-35 PEGMAs have shown good biocompat-
ibility,36,37 and their combination with a hydrophobic polymer
block such as poly(lactic acid-co-glycolic acid) (PLGA, which is
widely used clinically) has been shown to produce well-defined
and biocompatible nanoparticles.38,39 For encapsulation of bio-
polymers such as DNA, however, simple micelle formation with
these block materials is of limited utility, as the nucleic acids do
not easily partition to the interior phase. Water/oil/water emul-
sion systems as described above have been widely used to
encapsulate nucleic acids via their initially entrapped aqueous
regions, and block copolymer surfactants have been used to
enhance DNA incorporation into PLGA nanoparticles by this
route.40,41 PLGA-PEGMA block copolymers were therefore
obvious candidates to use for interfacial stabilization of w/o/w
emulsions, while the ability to tune block copolymer structure in
PLGA-PEGMA systems through controlled synthesis also, in
principle, allowed us further control of emulsion stability via this
route.
The synthetic strategy (Scheme 2) therefore combined two
different polymerization techniques, ring-opening polymeriza-
tion (ROP) and atom transfer radical polymerization (ATRP),
using an initial heterobifunctional initiator 2-bromo-2-methyl-
propionic acid 2-(2-hydroxyethyldisulfanyl)ethyl ester (1), fol-
lowed by well-established “click” chemistries as shown in Figure 1.
Synthesis of (2-Bromo-2-methylpropionic Acid 2-(2-Hy-
droxyethyldisulfanyl)ethyl Ester) (1). A heterobifunctional
unit was required to initiate the two polymerization routes, i.e.,
ROP for the synthesis of the core PLGA block and ATRP for the
synthesis of the polyPEGMA475 exterior block. We also required
the link between the two blocks to be cleavable, for which we
chose a disulfide bridge, as reduction of disulfides by intracellular
glutathione (GSH) has been previously described. Compound 1
was accordingly obtained by esterification of dithiodiethanol
with 2-bromo-2-methylpropionyl bromide. The structure of
the resultant 2-((2-hydroxyethyl) disulfanyl)ethyl 2-bromopro-
panoate was confirmed by IR, NMR, and mass spectrometry.
(ESI, Supporting Information Figure 1S).
Synthesis of PLGA-S-S-Br Polymers/macroinitiators. Ad-
dition of DL-lactide and glycolide to 2-((2-hydroxyethyl)disul-
fanyl)ethyl 2-bromopropanoate in the presence of tin(II) 2-
ethylhexanoate as catalyst resulted in ring-opening polymeriza-
tion of the monomers initiated from the hydroxyl terminus of 1.
The targeted molar mass for the PLGA block was ∼5000 Da, as is
apparent from gel permeation chromatography (GPC, ESI,
Supporting Information Table 1S, Figure 11S). The first block
was synthesized in the desired molar mass range and with
relatively low polydispersity compared to condensation-synthe-
1
sized PLGA. H NMR spectra (ESI, Suppoorting Information
Figure 2S) showed the expected resonances of the polymer
backbone at 5.1 ppm and 1.57 ppm (methine and methyl protons
of poly(lactic acid)) and at 4.8 ppm (methylene group of
poly(glycolic acid)).
Synthesis of PLGA-S-S-PEGMA475 Block Copolymer.
Atom transfer radical polymerization (ATRP) of PEGMA475
from the 2-bromo-2-methyl-propanoate-tipped PLGA was car-
ried out using the AGET ATRP method.42,43 The polymerization
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Figure 1. Transmission electron micrographs of different wt/wt% ratios of PLGA-S-S-PEGMA475-folate to PLGA-S-S-PLGA: (a) 10 wt % (DH = 187 (
14 nm), (b) 25 wt % (DH = 149 ( 15 nm), (c) 50 wt % (DH = 145 ( 29 nm), and (d) 75 wt % (DH = 172 ( 15 nm).
was stopped at low conversion to control polydispersity and
ensure a high degree of bromine end-functionality for subse-
quent derivatization. The polymerization of the second block was
confirmed by GPC (ESI, Supporting Information Figure 11S),
which showed a decrease in retention time of the PLGA-S-S-
PEGMA475 in comparison to PLGA-S-S-Br. The GPC traces
were unimodal, indicating successful synthesis of the second
block from the first block via the ATRP process. The molar mass
distribution of the PLGA-S-S-PEGMA475 block copolymer was
essentially the same as that of the precursor PLGA block which
per se demonstrated controlled polymerization for the second
component. The structure of the diblock copolymer was con-
firmed by 1H NMR spectra (ESI, Supporting Information Figure
3S), and the composition and ratios of each block in the co-
polymer were determined by relative peak integrals of PLGA and
PEGMA proton resonances (ESI).
Synthesis of PLGA-S-S-PLGA “Homodiblock” Copolymer.
A reducible “homodiblock” PLGA polymer, designed to act as a
bioresponsive core of the particles, was synthesized from dithio-
diethanol and DL-lactide and glycolide monomers via tin(II)
2-ethylhexanoate catalyzed ROP as before. The targeted molar
mass was again ∼5000 Da, but in this case, the copolymer grew
from both ends of the dithiodiethanol initiator. GPC traces (ESI,
Supporting Information Figure 11S) showed relatively good
control over polydispersity of the polymer, with the unimodal
feature of the GPC traces suggesting successful progressive chain
extension on either side of the initiator; the structure of the
diblock was also confirmed by spectroscopy (ESI).
Information Figure 11S). The 1H NMR spectrum (ESI, Support-
ing Information Figure 4S) was similar to that of the bromide-
terminated block copolymer with the exception of the methylene
residue at the terminus: this resonance shifted from 1.9 ppm in
the Br-terminated polymer to 1.7 ppm in the azide-derivative,
suggesting a complete substitution of the terminal bromine by
azide. FT-IR spectra confirmed the presence of azide function-
ality through a strong absorption band at ∼2110 cm-1. (ESI,
Supporting Information Figure 7S). The final step in the
synthetic procedure involved modification of the terminal azide
via copper-mediated “click” coupling with an acetylene-tipped
folic acid derivative, propargyl folate in DMF at room tempera-
ture. N,N0,N00,N000,N0000-Pentamethyldiethylenetriamine (PMDTA)
was employed to enhance reaction rate and increase end-
terminal modification. GPC again confirmed that no change in
molar mass occurred during the “click” reaction (ESI, Supporting
Information Figure 11S). FT-IR demonstrated the loss of azide
1
group absorption at 2110 cm-1 and H NMR showed the
presence of folate aromatic protons at δ = 6-8 ppm. UV
absorptions at 363 nm, representative of folate. provided further
support for the presence of azide functionality. The number of
folate residues per diblock copolymer of PLGA-S-S-PEGMA475
was determined to be ∼1 as expected.
Properties of synthesized polymers are listed in Table 1
Fabrication of Nanoparticles. The PLGA-S-S-PEGMA475
and PLGA-S-S-PLGA block copolymers were then used to
prepare blend matrix nanostructures as shown in Scheme 3.
The double emulsion technique was selected to encapsulate
DNA, as the formation of nanoparticles by this route occurs via
solvent diffusion, under conditions that are mild and avoid the
shear-induced degradation of DNA prevalent during conven-
tional nanoparticle formation. To accomplish this, an aqueous
solution of plasmid DNA was added into dichloromethane
(organic phase) containing PLGA-S-S-PLGA and folate-terminated
PLGA-S-S-PEGMA475. The primary emulsion needed only a
Synthesis of Azido- and Folate-Terminated PLGA-S-S-
PEGMA475 Diblock Copolymers. The bromine-terminated
PLGA-S-S-PEGMA475 diblock copolymer was transformed into
an azide-terminated diblock copolymer via nucleophilic substitu-
tion with sodium azide.44,45 The molar mass of the azide-
terminated PLGA-S-S-PEGMA475 was unchanged compared to
the bromo precursor as indicated by GPC (ESI, Supporting
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Table 1. Properties of Polymers and Diblock Copolymers Using ROP and ATRP Methods
polymer
Mn(th) ꢀ 103a
Mn ꢀ 103b
Mw ꢀ 103
Mw/Mn
PLGA-S-S-Br (P1)
4.9
9
4.9
8.4
8.1
9.4
9.4
4.2
4
6.5
10.2
9.7
12
1.3
1.2
1.2
1.3
1.3
1.2
1.2
PLGA-S-S-PEGMA475.Br (P2-a)
PLGA-S-S-PEGMA475.N3 (P2-b)
PLGA-S-S-PEGMA475.Folate (P3-a)
PLGA-S-S-PEGMA475-OH (P3-b)
PLGA-S-S-PLGA (P4)
9
9
9
12
4.9
4.9
5
PLGA-TEG-PLGA (P5)
4.7
a Theoretical, from monomer/initiator ratio. b From GPC (CHCl3), 30 °C, polystyrene standards).
Scheme 3. Preparation of Core-Shell Nanoparticles Using
Microscopy of nanoparticles (3% phosphotungstic acid nega-
tive staining) showed consistent spherical shape across the
compositional range (Figure 1). Particle sizes determined from
TEM micrographs were slightly larger than the hydrodynamic
diameters recorded by dynamic light scattering (DLS) and also
indicated a rather higher polydispersity of particle diameters. The
difference in overall diameters was most likely due to air-drying of
the nanoparticles on the copper grid leading to collapse of
hydrated polymerized PEGMA475 segments on the outer shells
and possibly some in situ aggregation. The apparent range of
particle sizes, i.e., the polydispersity of the sample by TEM, was
higher than observed in DLS. This may have been a consequence
of the relative insensitivity of light scattering techniques in
detecting smaller rather than larger particles: however, caution
is needed in interpreting TEM, as well as DLS, data owing to the
more rapid sedimentation of larger particles in TEM (Figure 1).
Encapsulation of Plasmid DNA. The combination of the
“homodiblock” PLGA-S-S-PLGA and the amphiphilic diblock PL-
GA-S-S-PEGMA475 polymers with the double emulsion method
resulted in efficient encapsulation of plasmid DNA (GWiz luciferase).
The extent of incorporation was 80% of the DNA in solution, co-
rresponding to a theoretical loading of 0.4% with respect to the mass
of polymer in the emulsion (∼4 μg DNA per mg of polymer).
Recovery of DNA from the nanoparticles following treatment of the
dried copolymers with 0.5% N NaOH indicated that 3.2 μg.mg of
plasmid DNA had been encapsulated in the particles as determined by
UV spectroscopy and comparison to standard absorption values at
260 nm. (ESI, Supporting Information Table 2S).
In Vitro Release Profiles of Plasmid DNA from Nanopar-
ticles. The release profiles of DNA from the nanoparticles were
investigated under conditions designed to mimic reducing and
nonreducing environments in the cytosol. Copolymer particles
in 50:50 molar ratios of PLGA-S-S-PEGMA475 to either PLGA-
S-S-PLGA (reducible) or PLGA-TEG-PLGA (nonreducible)
were loaded with the same amounts of plasmid DNA (3.2 μg.mg
polymer) and incubated in PBS pH 7.4 in the presence or
absence of 5 mM glutathione (GSH). As apparent from Figure 2,
the disulfide core based nanoparticles showed 85% release of
DNA over 16 h in the presence of 5 mM GSH compared to 40%
release in PBS buffer alone. Furthermore, there was no significant
difference between release of DNA from nanoparticles contain-
ing the nonreducible PLGA-TEG-PLGA core system in the
presence and absence of 5 mM GSH (45% and 46% release,
respectively). An initial burst release was observed for all the
nanoparticle systems over a 2 h period as has been reported
before for PLGA-based nano- and microparticles,46-48 but only
for the reducible PLGA-S-S-PLGA core materials in the presence
of GSH was the rapid release continued over the full 16 h period
monitored.
the W/O/W Emulsion Techniquea
a In (a), DNA in water is added into dichloromethane containing PLGA-
S-S-PEGMA475 polymers. Addition of this primary emulsion to ethanol
(b) results in collapse of the hydrophobic core as the CH2Cl2 partitions
into the bulk phase. Solvent evaporation (c) yields the final nano-
particles.
mild vortex to form, which was most likely a result of the
amphiphilic nature of the folate-terminated PLGA-S-S-PEG-
MA475. Addition of this emulsion to a more polar phase
(ethanol) resulted in the immediate formation of nanoparticles
as dichloromethane diffused from the inner oil phase to be
replaced by ethanol, a nonsolvent for the PLGA block. The final
nanoparticles obtained after solvent evaporation were of low
polydispersity as determined by TEM size distributions and
DLS and net negative charge. Several batches of nanoparticles
were prepared by this route using different ratios of PLGA-S-
S-PLGA to folate-terminated PLGA-S-S-PEGMA475 diblock
copolymer.
As apparent from Table 2, increasing the PEGMA475-folate/
PLGA-S-S-PLGA ratio (NPS1-5) decreased both particle size
and zeta potential. This was likely due to the greater interfacial
stabilization during primary emulsion formation as the PEG-
MA475 content was increased, and the shielding of PLGA-
carboxyl groups as the PEGMA475-folate “corona” was enlarged.
Particle sizes were similar for NPS4 and NPS6, which contained
PLGA-S-S-PEGMA475-folate/PLGA-S-S-PLGA (50:50 mol %)
and PLGA-S-S-PEGMA475-OH/PLGA-S-S-PLGA (50:50 mol %),
respectively, at ∼110 nm, when DNA was not encapsulated.
Similarly, NPS 8 and NPS 9, which were of the same molar ratio
as PLGA-S-S-PEGMA475-folate or PLGA-S-S-PEGMA475-OH
but with a nonreducible PLGA-TEG-PLGA in place of PLGA-S-
S-PLGA, were formed as ∼110 nm particles in the absence of
DNA. However, for NPS 4, 8, and 9, the encapsulation of DNA
decreased both particle size and zeta potential. These differences
in size and charge of the particles containing DNA indicate that
the nucleic acids may also have played a role in stabilizing the
initial o/w interface in the primary emulsion.
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Table 2. Mean Particle Size and Zeta Potential of the Different Blank and Plasmid-Loaded PLGA-S-S-PLGA and PLGA-S-S-
PEGMA475-Folate Nanoparticles (mean ( S.D., n = 3)
batch
ratioa
blank mean size (nm)b
plasmid-loaded mean size (nm)
blank zeta potential (mV)
plasmid loaded zeta potential (mV)
NPS1
NPS2
NPS3
NPS4
NPS5
NPS6
NPS7
NPS8
NPS9
95:5c
148 ( 1
145 ( 2
128 ( 4
108 ( 4
95 ( 3
......
-15.6 ( 0.3
-14.1 ( 0.1
-6.9 ( 0.4
-7.2 ( 0.2
-4.5 ( 0.4
-13.3 ( 0.3
-15.1 ( 0.2
-8.5 ( 0.1
-13.1 ( 0.1
......
......
90:10c
75:25c
50:50d
25:75d
50:50d
25:75d
50:50e
50:50f
......
......
......
96 ( 1
......
-9.9 ( 0.1
......
110 ( 5
139 ( 3
111 ( 1
108 ( 1
......
......
......
......
98 ( 3
102 ( 1
-9.5 ( 0.1
-13.6 ( 0.1
a Wt/wt % ratio of PLGA-S-S-PLGA/PLGA-S-S-PEGMA475-folate. b Mean diameter based on relative mass distribution. c Mean diameter derived from
P3a (folate-terminated). d Mean diameter derived from P3b (hydroxyl-terminated). e Derived as in c, except P5 (PLGA-TEG-PLGA) was used instead
of P4. f Derived as in d, except P5 (PLGA-TEG-PLGA) was used instead of P4.
Figure 2. Cumulative release (%) of plasmid DNA from nanoparticles.
Comparison of nanoparticles with reducible cores (PLGA-S-S-PLGA)
compared to nonreducible cores (PLGA-TEG-PLGA) under nonredu-
cing conditions (PBS, pH 7.4) and cytosol-mimicking reducing condi-
tions (5 mM glutathione, GSH) shows rapid and higher cumulative
release for the reducible core polymers under reducing conditions.
Figure 3. Folate-mediated cellular uptake of nanoparticles by Calu-3
cell lines. The uptake of particles (expressed as the equivalent concen-
tration of particles in suspension in mg.mL-1) is shown on the y axis,
with increasing functionalization of nanoparticles with folate (NPS
1-5) shown with horizontal-striped bars on the x axis. For each
functionalized particle, the uptake when 1 mM of folic acid was present
in the cell suspension is shown (vertical striped bars), while the uptake of
a hydroxyl-tipped PLGA-PEGMA475 polymer is shown as a comparison
(gray bar).
Uptake of Nanoparticles in Calu-3 Cell Line. The lung
cancer cell line Calu-3 was chosen as a suitable and clinically
relevant in vitro test of nanoparticle gene delivery efficacy.
Recent studies have indicated that pemetrexed, a folate anti-
metabolite, is a potent and effective therapeutic in the treatment
of lung cancers,49 a finding that suggests that folate-receptor
(FR) mediated targeting50 might be highly suitable for lung
cancer cell lines. Accordingly, we generated Calu-3 monolayers
and incubated them with nanoparticle suspensions when the cells
were confluent and linked via tight junctions as shown by trans-
epithelial electrical resistance (TEER) measurements (data not
shown). The nanoparticles in the uptake assay contained pro-
gressively increasing molar ratios of PLGA-S-S-PEGMA475-fo-
late relative to the PLGA-S-S-PLGA and PLGA-TEG-PLGA
“cores”, while “nonfunctional” PLGA-S-S-PEGMA475-OH/
PLGA-S-S-PLGA nanoparticles were used as a control for the
PEGMA-folate surface. As a positive control, a second set of
assays was carried out, this time in the presence of 1 mM free folic
acid, to block any specific FR-mediated uptake. The results of
these assays are shown in Figure 3.
Cell uptake of nanoparticles increased with increasing
amounts of the folate-tipped block copolymers in the matrix,
although there was only a marginal difference between uptake of
the 75% folate-tipped particles (NPS5) compared to 50%
(NPS4). Uptake of the hydroxyl-terminated particles (NPS9)
was significantly lower (p = 0.002) than that of the folate-
terminated systems. In all cases, addition of free folic acid
suppressed uptake, consistent with a folate-receptor mediated
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Figure 5. Cell metabolic activities as a marker of viability of Calu-3 cells
as evaluated by the MTS assay following incubation with NPS1 and
NPS5 particles compared to TritonX (positive control) and HBSS
(negative control). Activity levels are represented as a percentage with
respect to the level obtained for the negative control (buffer solution
HBSS). Error bars show standard deviations from the mean.
Figure 4. Transfection assays in Calu-3 monolayers with nanoparticles
with/without cell-targeting and bioreducible functionality. RLU =
relative light units, a measure of luciferase expression. Highest transgene
expression in Calu-3 cells was exhibited by the reducible folate-targeted
nanoparticles (NPS 4, vertical striped bar) and overall expression was
further enhanced by addition of exogenous GSH (NPS 4, gradient gray
bar). Error bars show standard deviations from the mean.
enhancing the performance of the modular polymeric gene
delivery system.
Cell metabolic activity assays following the transfection ex-
periments indicated that the polymer nanoparticles were well-
tolerated by Calu-3 cells. At concentrations of up to 4 mg.mL-1
,
nanoparticles NPS 1 and NPS 5 did not reduce cell metabolic
activity in Calu-3 cells, as measured by the MTS assay, by greater
than 20% (Figure 5), with the 5% folate-functionalized particles
(NPS1) resulting in reductions of ∼5%.
uptake mechanism, rather than that due to the small variations
((10%) in particle size over the nanoparticle samples.
These data indicate the nanoparticles were of overall low
toxicity. Taken together, the specificity of the uptake in the
absence and presence of free folic acid and/or “nonfunctional”
OH-terminal particles, combined with the enhancements ob-
served for transgene expression using folate-functional particles
with bioreducible disulfide cores and the lack of toxicity, indicate
that the key design goals had been attained.
The key experiments, bringing together all the design compo-
nents of the system, i.e., DNA incorporation, core-shell archi-
tecture with surface-displayed cell targeting ligands, and
intracellularly reducible core for enhanced release, involved
evaluation of transgene expression in Calu-3 monolayers, utiliz-
ing the GWiz luciferase plasmid. Three nanoparticle systems
were chosen: NPS9 with a nonreducible core and a PEGMA475
shell with no folate termini, NPS8 with a PEGMA475-folate
surface but a nonreducible core, and NPS4 with both PEGMA475
folate surfaces and the PLGA-S-S-PLGA reducible core. As a
further probe of activity, transfections with NPS4 were also carried
out with addition of glutathione monoethyl ester (GSH-MEE),
which is converted into active GSH intracellularly, 4 h after initial
incubation of the cells with the particles. Light production due to
expressed luciferase in Calu-3 is shown in Figure 4.
’ DISCUSSION
Our overall aim in this study was to establish the chemistries
necessary for incorporating a wide range of potential biother-
apeutic agents, such as proteins and nucleic acids, inside protec-
tive drug delivery vehicles that should be biocompatible, easily
functionalizable for cell targeting, and bioresponsive. An impor-
tant part of the strategy was that the components should be
modular, so that different parts could be combined together or
replaced, depending on a particular desired application. We
chose a PLGA core because this polymer has been extensively
used clinically, and the degradation rate can be tuned by
comonomer content and degree of crystallinity.51 For some gene
delivery applications such as cancer therapies, rapid release is
usually required, but for a slower degradation profile as might be
considered for longer-term tissue engineering/regeneration ap-
plications, poly(caprolactone) (PCL) or other polyesters could
be used instead of PLGA. For more rapid release, PGA or other
more hydrolytically sensitive polyesters could be employed as the
inner block. For all these modifications, the initial bifunctional
initiator 2-bromo-2-methyl-propionic acid 2-(2-hydroxyethyldi-
sulfanyl)ethyl ester) (1) could be employed with variation in
monomers as appropriate. In addition, the use of the disulfide
Overall transgene expression was higher for the bioreducible
and folate-targeted nanoparticles, with the highest light produc-
tion for cells incubated with NPS4 with addition of GSH-MEE.
Transfection increased 5-fold for the folate-functional NPs (NPS
2) compared to the nontargeted nonreducible NPs (NPS 9) and
10-fold by incorporation of the additional reducible component
in the cores (NPS 4). Boosting the intracellular GSH resulted in
an overall increase of ∼4-fold (NPS 4 þ GSH) in luciferase
expression compared to that without exogenous GSH, and 42-
fold (NPS 4þGSH) compared to the nonreducible nontargeted
NPs (NPS9). These increases in expression of the transgene
indicated that DNA was released from the nanoparticles in a form
able to be transcribed and translated, and showed that the
introduction of the targeted shell layer and bioreducible core
components added synergistically, as well as individually, to
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linker allows further control of release rate,52,53 as a variety of
disulfide containing blocks that are themselves rapidly cleavable
could be used to grow PLGA, PGA, or PCL cores. Examples exist
of cysteine-terminated peptides, linked together by disulfides to
form polymers. These have been shown to condense or encap-
sulate DNA during transport and to release the nucleic acid via a
cytosolic reduction mechanism:54,55 such peptides could also be
incorporated into the core of our polymers. The outer compo-
nent, poly(PEGMA), enables additional flexibility in delivery
vehicle design. We utilized a PEGMA475 homopolymer grown by
ATRP, but copolymers with other PEG-methacrylates offer the
possibilities of tumor targeting through a thermally responsive
phase transition,34,35 while an arsenal of targeting functionality in
addition to folate is readily accessible through the wealth of
“click” chemistries reported in the literature.56 Recent data have
indicated biocompatibility of PEGMA-based polymers grown by
controlled radical routes36 providing that catalyst residues are
effectively removed,37 and thus, it is likely that a platform of
therapeutic carrier materials could be produced by the route we
demonstrate here. When combined with the nanoprecipitation
method of particle formation, this approach offers advantages
over other gene delivery vector syntheses, as it allows systematic
and easily variable placement of functional components in order
to provide a stable but readily degradable core, a cleavable steric
shield, and efficient introduction of targeting ligands at the site
they are needed, i.e., the outer surface.
We illustrated the efficacy of our modular polymers using
Calu-3 cells, as a clinically important differentiated lung cancer
cell line used before as a model for nasal and diseased airway
epithelia. Calu-3 cells were also chosen because they have
previously been shown to be hard to transfect.23 Overall levels
of transgene expression were very similar (∼1-200 RLU/mg
protein) in luciferase assays compared with PEI polyplexes, and
higher than for the cationic lipid DOTAP, based on data reported
in the prior studies by Florea et al.23 However, the PLGA-
PEGMA-folate nanoparticles in our study were as effective as PEI
without inducing the same decreases in metabolic activity.23
Calu-3 cells grown under the same conditions as we used, i.e., at
the air-liquid interface, have been reported to secrete mucus,
and this has been shown to be an effective barrier to PEI-
mediated transfection in vitro, with much lower levels of
transfection in Calu-3 compared to the undifferentiated, non-
mucus-producing A549 cell line.57 However, the relative levels of
transfection with different PEI-based polyplexes in Calu-3 levels
were much better predictors for in vivo activity than transfection
levels in A549 cells. While the ultimate test for our new modular
polymers is of course in vivo, the levels of transgene expression
using our approach in vitro are promising and should follow the
patterns we obtained in Calu-3 in a manner analogous to the in
vitro/in vivo correlation observed by Nguyen et al.57 Impor-
tantly, the colloidal stability of polyPEGMA-PLGA nanoparticles
with encapsulated DNA (no aggregation after 16 h) suggests that
our materials should be easier to formulate for in vivo use than
conventional PEI/DNA complexes, which readily aggregate on
standing (∼2 h).57 In addition, the synthesis of the outer
polyPEGMA475 layers through ATRP routes enables better
control of steric shield layer thickness and end-terminal func-
tionality than is possible through end-functionalization of lipo-
somes by preformed PEGylation reagents. This in turn should
lead to greater control over variables important in determining
pharmacokinetics and targeting in vivo. For ultimate use in the
clinic, it is necessary to ensure high transfection efficiency but
without compromising safety, and the modular nature of these
systems could be adapted further to include components that
exploit biological mechanisms to enhance transfection. Thus, the
use of endosomal buffering polymers and nuclear-localizing
sequences as parts of the outer polymer shell could be envisaged,
as numerous literature reports have indicated their efficacy in
other gene delivery systems.58-61 Experiments to evaluate the in
vivo efficacy of the “first-generation” polyPEGMA-PLGA nano-
particles we report here are now ongoing in our laboratories and
will be reported in a future manuscript.
The synthetic method we have described has one final
advantage in that the use of the double emulsion route allows
blending of different polymers to tune the degree of core-vs-shell
components, and to encapsulate a variety and/or combination of
potent yet delicate biomacromolecules with high efficiency. One
could thus envisage nucleic acids for gene delivery and growth
factors for tissue development incorporated in particles made by
essentially the same methods but with different core components
to allow phased release of actives.
’ CONCLUSIONS
Modular block copolymers were assembled via ROP, ATRP,
and “click” chemistry routes to generate biocompatible, cell-
targeted, and bioresponsive gene delivery vehicles. Double
emulsion techniques were used to encapsulate DNA with high
efficiency, and uptake of the resultant nanoparticles was shown to
take place in Calu-3 cells via a specific folate-receptor mechanism.
The transgene expression studies demonstrated that selective
receptor-mediated and intracellularly triggered drug delivery
took place, and the ability to do this in the clinically important
Calu-3 cell line offers promise for future therapeutic applications.
’ ASSOCIATED CONTENT
Supporting Information. 1HNMR, GPC chromato-
S
b
graph, and FT-IR of the polymers, standard calibration curves
of GWIZ luciferase plasmid and folic acid. This material is
available free of charge via the Internet at http://pubs.acs.org.
’ AUTHOR INFORMATION
Corresponding Author
*School of Pharmacy, University of Nottingham, Nottingham,
NG7 2RD, U.K. Tel þ44 115 846 7678; Fax þ44 115 951 5102;
E-mail cameron.alexander@nottingham.ac.uk.
’ ACKNOWLEDGMENT
We thank Suleimaniyeh University, Iraq (scholarship for
Aram Omer Saeed), the Egyptian Government. (scholarship to
Mahmoud Soliman), the Engineering and Physical Sciences
Research Council (EPSRC Grant EP/H005625/1) and the
University of Nottingham for financial support. Christine Grain-
ger-Boultby, School of Pharmacy, University of Nottingham is
thanked for technical support and Drs. Giuseppe Mantovani and
Martin Garnett for many helpful discussions.
’ REFERENCES
(1) Lee, K. D. (2005) New fashions in molecular pharmaceutics:
Nanoscience and nanomedicine. Does size matter in delivery? Mol.
Pharmaceutics 2, 169–169.
166
dx.doi.org/10.1021/bc100149g |Bioconjugate Chem. 2011, 22, 156–168

Bioconjugate Chemistry
ARTICLE
(2) Wong, S. Y., Pelet, J. M., and Putnam, D. (2007) Polymer
systems for gene delivery-past, present, and future. Prog. Polym. Sci. 32,
799–837.
(3) Sawant, R. M., Hurley, J. P., Salmaso, S., Kale, A., Tolcheva, E.,
Levchenko, T. S., and Torchilin, V. P. (2006) ”SMART” drug delivery
systems: Double-targeted pH-responsive pharmaceutical nanocarriers.
Bioconjugate Chem. 17, 943–949.
(4) Barua, S., Joshi, A., Banerjee, A., Matthews, D., Sharfstein, S. T.,
Cramer, S. M., Kane, R. S., and Rege, K. (2009) Parallel synthesis and
screening of polymers for nonviral gene delivery. Mol. Pharmaceutics 6,
86–97.
(5) Cavalieri, F., Postma, A., Lee, L., and Caruso, F. (2009) Assembly
and functionalization of DNA-polymer microcapsules. ACS Nano 3,
234–240.
(6) Guillot-Nieckowski, M., Joester, D., Stohr, M., Losson, M.,
Adrian, M., Wagner, B., Kansy, M., Heinzelmann, H., Pugin, R.,
Diederich, F., and Gallani, J. L. (2007) Self-assembly, DNA complexa-
tion, and pH response of amphiphilic dendrimers for gene transfection.
Langmuir 23, 737–746.
(7) Zhang, L., Nguyen, T. L. U., Bernard, J., Davis, T. P., Barner-Kowollik,
C., and Stenzel, M. H. (2007) Shell-cross-lihked micelles containing cationic
polymers synthesized via the RAFT process: Toward a more biocompatible
gene delivery system. Biomacromolecules 8, 2890–2901.
(8) Bromberg, L., Deshmukh, S., Temchenko, M., Iourtchenko, L.,
Alakhov, V., Alvarez-Lorenzo, C., Barreiro-Iglesias, R., Concheiro, A.,
and Hatton, T. A. (2005) Polycationic block copolymers of poly-
(ethylene oxide) and poly(propylene oxide) for cell transfection.
Bioconjugate Chem. 16, 626–633.
self-immolative dendrimers with poly(ethylene glycol) via click chem-
istry. Bioconjugate Chem. 17, 1432–1440.
(19) Lee, S., Greenwald, R. B., Mcguire, J., Yang, K., and Shi, C.
(2001) Drug delivery systems employing 1,6-elimination: releasable
poly(ethylene glycol) conjugates of proteins. Bioconjugate Chem. 12,
163–169.
(20) Van Butsele, K., Jerome, R., and Jerome, C. (2007) Functional
amphiphilic and biodegradable copolymers for intravenous vectorisa-
tion. Polymer 48, 7431–7443.
(21) Low, P. S., Henne, W. A., and Doorneweerd, D. D. (2008)
Discovery and development of folic-acid-based receptor targeting for
imaging and therapy of cancer and inflammatory diseases. Acc. Chem.
Res. 41, 120–129.
(22) De, P., Gondi, S. R., and Sumerlin, B. S. (2008) Folate-
conjugated thermoresponsive block copolymers: Highly efficient con-
jugation and solution self-assembly. Biomacromolecules 9, 1064–1070.
(23) Florea, B. I., Meaney, C., Junginger, H. E., and Borchard, G.
(2002) Transfection efficiency and toxicity of polyethylenimine in
differentiated Calu-3 and nondifferentiated COS-1 cell cultures. AAPS
PharmSci 4, 1–11.
(24) Zhu, Y., Chidekel, A., and Shaffer, T. H. (2010) Cultured
Human Airway Epithelial Cells (Calu-3): A Model of Human Respira-
tory Function, Structure, and Inflammatory Responses. Critical Care
Research and Practice, ID 394578, 8 pages.
(25) Zhang, K., Rossin, R., Hagooly, A., Chen, Z. Y., Welch, M. J.,
and Wooley, K. L. (2008) Folate-mediated cell uptake of shell-cross-
linked spheres and cylinders. J. Polym. Sci., Part A: Polym. Chem. 46,
7578–7583.
(9) Bartlett, D. W., and Davis, M. E. (2007) Physicochemical and
biological characterization of targeted, nucleic acid-containing nanopar-
ticles. Bioconjugate Chem. 18, 456–468.
(26) Pan, D., Turner, J. L., and Wooley, K. L. (2003) Folic acid-
conjugated nanostructured materials designed for cancer cell targeting.
Chem. Commun. 2400–2401.
(10) Park, J. S., Akiyama, Y., Yamasaki, Y., and Kataoka, K. (2007)
Preparation and characterization of polyion complex micelles with a
novel thermosensitive poly(2-isopropyl-2-oxazoline) shell via the com-
plexation of oppositely charged block ionomers. Langmuir 23, 138–146.
(11) Miyata, K., Fukushima, S., Nishiyama, N., Yamasaki, Y., and
Kataoka, K. (2007) PEG-based block catiomers possessing DNA
anchoring and endosomal escaping functions to form polyplex micelles
with improved stability and high transfection efficacy. J. Controlled
Release 122, 252–260.
(27) You, Y. Z., Zhou, Q. H., Manickam, D. S., Wan, L., Mao, G. Z.,
and Oupicky, D. (2007) Dually responsive multiblock copolymers via
reversible addition-fragmentation chain transfer polymerization: Synth-
esis of temperature- and redox-responsive copolymers of poly(N-
isopropylacrylamide) and poly(2-(dimethylamino)ethyl methacrylate).
Macromolecules 40, 8617–8624.
(28) Mckenzie, D. L., Smiley, E., Kwok, K. Y., and Rice, K. G. (2000)
Low molecular weight disulfide cross-linking peptides as nonviral gene
delivery carriers. Bioconjugate Chem. 11, 901–909.
(12) Fukushima, S., Miyata, K., Nishiyama, N., Kanayama, N.,
Yamasaki, Y., and Kataoka, K. (2005) PEGylated polyplex micelles from
triblock catiomers with spatially ordered layering of condensed pDNA
and buffering units for enhanced intracellular gene delivery. J. Am. Chem.
Soc. 127, 2810–2811.
(13) Bae, Y., Nishiyama, N., and Kataoka, K. (2007) In vivo
antitumor activity of the folate-conjugated pH-Sensitive polymeric
micelle selectively releasing adriamycin in the intracellular acidic com-
partments. Bioconjugate Chem. 18, 1131–1139.
(14) Bae, Y., Nishiyama, N., Fukushima, S., Koyama, H., Yasuhiro,
M., and Kataoka, K. (2005) Preparation and biological characterization
of polymeric micelle drug carriers with intracellular pH-triggered drug
release property: Tumor permeability, controlled subcellular drug
distribution, and enhanced in vivo antitumor efficacy. Bioconjugate
Chem. 16, 122–130.
(15) Bae, Y., Jang, W. D., Nishiyama, N., Fukushima, S., and Kataoka,
K. (2005) Multifunctional polymeric micelles with folate-mediated
cancer cell targeting and pH-triggered drug releasing properties for
active intracellular drug delivery. Mol. Biosyst. 1, 242–250.
(16) Shuai, X. T., Merdan, T., Unger, F., Wittmar, M., and Kissel, T.
(2003) Novel biodegradable ternary copolymers hy-PEI-g-PCL-b-PEG:
Synthesis, characterization, and potential as efficient nonviral gene
delivery vectors. Macromolecules 36, 5751–5759.
(29) Han, G., Chari, N. S., Verma, A., Hong, R., Martin, C. T., and
Rotello, V. M. (2005) Controlled recovery of the transcription of
nanoparticle-bound DNA by intracellular concentrations of glutathione.
Bioconjugate Chem. 16, 1356-1359.
(30) Licciardi, M., Tang, Y., Billingham, N. C., and Armes, S. P. (2005)
Synthesis of novel folic acid-functionalized biocompatible block copolymers
by atom transfer radical polymerization for gene delivery and encapsulation of
hydrophobic drugs. Biomacromolecules 6, 1085–1096.
(31) Oh, J. K., Siegwart, D. J., Lee, H. I., Sherwood, G., Peteanu, L.,
Hollinger, J. O., Kataoka, K., and Matyjaszewski, K. (2007) Biodegrad-
able nanogels prepared by atom transfer radical polymerization as
potential drug delivery carriers: Synthesis, biodegradation, in vitro
release, and bioconjugation. J. Am. Chem. Soc. 129, 5939–5945.
(32) Lutz, J. F., Neugebauer, D., and Matyjaszewski, K. (2003)
Stereoblock copolymers and tacticity control in controlled/living radical
polymerization. J. Am. Chem. Soc. 125, 6986–6993.
(33) Lutz, J. F. (2008) Polymerization of oligo(ethylene glycol)
(meth)acrylates: Toward new generations of smart biocompatible
materials. J. Polym. Sci., Part A: Polym. Chem. 46, 3459–3470.
(34) Lutz, J. F., Andrieu, J., Uzgun, S., Rudolph, C., and Agarwal, S.
(2007) Biocompatible, thermoresponsive, and biodegradable: Simple
preparation of “all-in-one” biorelevant polymers. Macromolecules 40,
8540–8543.
(17) Tomlinson, R., Heller, J., Brocchini, S., and Duncan, R. (2003)
Polyacetal-doxorubicin conjugates designed for pH-dependent degrada-
tion. Bioconjugate Chem. 14, 1096–1106.
(18) Gopin, A., Ebner, S., Attali, B., and Shabat, D. (2006) Enzymatic
activation of second-generation dendritic prodrugs: Conjugation of
(35) Magnusson, J. P., Khan, A., Pasparakis, G., Saeed, A. O., Wang,
W., and Alexander, C. (2008) Ion-sensitive “isothermal” responsive
polymers prepared in water. J. Am. Chem. Soc. 130, 10852–10853.
(36) Ryan, S. M., Wang, X. X., Mantovani, G., Sayers, C. T.,
Haddleton, D. M., and Brayden, D. J. (2009) Conjugation of salmon
167
dx.doi.org/10.1021/bc100149g |Bioconjugate Chem. 2011, 22, 156–168

Bioconjugate Chemistry
ARTICLE
calcitonin to a combed-shaped end functionalized poly(poly(ethylene
glycol) methyl ether methacrylate) yields a bioactive stable conjugate. J.
Controlled Release 135, 51–59.
(37) Chang, C. W., Bays, E., Tao, L., Alconcel, S. N. S., and Maynard,
H. D. (2009) Differences in cytotoxicity of poly(PEGA)s synthesized by
reversible addition-fragmentation chain transfer polymerization. Chem.
Commun. 3580–3582.
(38) Saeed, A. O., Dey, S., Howdle, S. M., Thurecht, K. J., and
Alexander, C. (2009) One-pot controlled synthesis of biodegradable and
biocompatible co-polymer micelles. J. Mater. Chem. 19, 4529–4535.
(39) Abulateefeh, S. R., Saeed, A. O., Aylott, J. W., Chan, W. C.,
Garnett, M. C., Saunders, B. R., and Alexander, C. (2009) Facile
synthesis of responsive nanoparticles with reversible, tuneable and rapid
thermal transitions from biocompatible constituents. Chem. Commun..
(40) Gu, W. W., Xu, Z. H., Gao, Y., Chen, L. L., and Li, Y. P. (2006)
Transferrin-mediated PEGylated nanoparticles for delivery of DNA/
PLL. Nanotechnology 17, 4148–4155.
(2008) Combination dual responsive polypeptide vectors for enhanced
gene delivery. Mol. BioSyst. 4, 741–745.
(56) Lutz, J. F., Borner, H. G., and Weichenhan, K. (2006) Combin-
ing ATRP and “click” chemistry: a promising platform toward functional
biocompatible polymers and polymer bioconjugates. Macromolecules 39,
6376–6383.
(57) Nguyen, J., Xie, X., Neu, M., Dumitrascu, R., Reul, R., Sitterberg,
J., Bakowsky, U., Schermuly, R., Fink, L., Schmehl, T., Gessler, T., Seeger,
W., and Kissel, T. (2008) Effects of cell-penetrating peptides and pegyla-
tion on transfection efficiency of polyethylenimine in mouse lungs. J. Gene
Med. 10, 1236–1246.
(58) Stevenson, M., Ramos-Perez, V., Singh, S., Soliman, M., Preece,
J. A., Briggs, S. S., Read, M. L., and Seymour, L. W. (2008) Delivery of
siRNA mediated by histidine-containing reducible polycations. J. Con-
trolled Release 130, 46–56.
(59) Kang, H. C., and Bae, Y. H. (2007) pH-tunable endosomolytic
oligomers for enhanced nucleic acid delivery. Adv. Funct. Mater. 17,
1263–1272.
(60) Oishi, M., Nagatsugi, F., Sasaki, S., Nagasaki, Y., and Kataoka, K.
(2005) Smart polyion complex micelles for targeted intracellular
delivery of PEGylated antisense oligonucleotides containing acid-labile
linkages. ChemBioChem 6, 718–725.
(61) Knorr, V., Russ, V., Allmendinger, L., Ogris, M., and Wagner, E.
(2008) Acetal linked oligoethylenimines for use as pH-sensitive gene
carriers. Bioconjugate Chem. 19, 1625–1634.
(41) Diez, S., and de Ilarduya, C. T. (2006) Versatility of biodegrad-
able poly(D,L-lactic-co-glycolic acid) microspheres for plasmid DNA
delivery. Eur. J. Pharm. Biopharm. 63, 188–197.
(42) Tang, H. D., Radosz, M., and Shen, Y. Q. (2006) CuBr2/N,N,
0
0
N ,N -tetra[(2-pyridal)methyl] ethylenediamine/tertiary amine as a
highly active and versatile catalyst for atom-transfer radical polymeriza-
tion via activator generated by electron transfer. Macromol. Rapid
Commun. 27, 1127–1131.
(43) Jakubowski, W., and Matyjaszewski, K. (2005) Activator gen-
erated by electron transfer for atom transfer radical polymerization.
Macromolecules 38, 4139–4146.
(44) Lutz, J. F., and Zarafshani, Z. (2008) Efficient construction of
therapeutics, bioconjugates, biomaterials and bioactive surfaces using
azide-alkyne “click” chemistry. Adv. Drug Delivery Rev. 60, 958–970.
(45) Mantovani, G., Ladmiral, V., Tao, L., and Haddleton, D. M.
(2005) One-pot tandem living radical polymerisation-Huisgens cycload-
dition process (’’click’’) catalysed by N-alkyl-2-pyridylmethanimine/
Cu(I)Br complexes. Chem. Commun. 2089–2091.
(46) Tinsley-Bown, A. M., Fretwell, R., Dowsett, A. B., Davis, S. L.,
and Farrar, G. H. (2000) Formulation of poly(-lactic-co-glycolic acid)
microparticles for rapid plasmid DNA delivery. J. Controlled Release 66,
229–241.
(47) Wang, D. Q., Robinson, D. R., Kwon, G. S., and Samuel, J.
(1999) Encapsulation of plasmid DNA in biodegradable poly(D,L-
lactic-co-glycolic acid) microspheres as a novel approach for immuno-
gene delivery. J. Controlled Release 57, 9–18.
(48) Luo, D., Woodrow-Mumford, K., Belcheva, N., and Saltzman,
W. M. (1999) Controlled DNA delivery systems. Pharm. Res. 16, 1300–
1308.
(49) Bunn, P. A., and Thatcher, N. (2008) Systemic treatment for
advanced (Stage IIb/IV) non-small cell lung cancer: More treatment
options: More things to consider: Introduction. Oncologist 13, 1–4.
(50) Ross, J. F., Chaudhuri, P. K., and Ratnam, M. (1994) Differ-
ential regulation of folate receptor isoforms in normal and malignant
tissues in vivo and in established cell lines. Physiologic and clinical
implications. Cancer 73, 2432–2443.
(51) Bala, I., Hariharan, S., and Kumar, M. (2004) PLGA nanopar-
ticles in drug delivery: The state of the art. Crit. Rev. Ther. Drug Carrier
Syst. 21, 387–422.
(52) Li, Y. T., and Armes, S. P. (2005) Synthesis and chemical degrada-
tion of branched vinyl polymers prepared via ATRP: Use of a cleavable
disulfide-based branching agent. Macromolecules 38, 8155–8162.
(53) Hong, R., Han, G., Fernandez, J. M., Kim, B. J., Forbes, N. S.,
and Rotello, V. M. (2006) Glutathione-mediated delivery and release
using monolayer protected nanoparticle carriers. J. Am. Chem. Soc. 128,
1078–1079.
(54) Wan, L., Manickam, D. S., Oupicky, D., and Mao, G. Z. (2008)
DNA release dynamics from reducible polyplexes by atomic force
microscopy. Langmuir 24, 12474–12482.
(55) Nasanit, R., Iqbal, P., Soliman, M., Spencer, N., Allen, S., Davies,
M. C., Briggs, S. S., Seymour, L. W., Preece, J. A., and Alexander, C.
168
dx.doi.org/10.1021/bc100149g |Bioconjugate Chem. 2011, 22, 156–168