The controlled intravenous delivery of drugs using PEG-coated sterically [606321]

The controlled intravenous delivery of drugs using PEG-coated sterically
stabilized nanospheres ☆
R. Grefa,⁎, A. Dombb, P. Quelleca, T. Blunkc, R.H. Müllerd, J.M. Verbavatze, R. Langerf
aLaboratoire de Chimie-Physique Macromoléculaire, URA CNRS 494, ENSIC, 1, Rue Grandville, BP 451, 54001 Nancy Cedex, France
bDepartment of Pharmaceutical Chemistry, Hebrew University, 91120 Jerusalem, Israel
cDepartment of Pharmaceutics and Biopharmaceutics, Christian-Albrechts-University Kiel, Gutenbergstr, 76 –78, 24118 Kiel, Germany
dDepartment of Pharmaceutics, Biopharmaceutics and Biotechnology, Free University of Berlin, Kelchstr, 31, 12169 Berlin, Germany
eService de Biologie Cellulaire, CEA Saclay, 91191 Gif sur Yvette, France
fDepartment of Chemical Engineering, MIT, E25-342, 45 Carlton St., Cambridge, MA 02139, USA
abstract article info
Article history:
Accepted 13 April 1995
Available online 13 September 2012
Keywords:Long-circulating nanoparticlesBiodegradable polymersPolyethylene glycol
Hydrophilic coating
Reduced liver accumulationIntravenous drug administrationInjectable blood persistent particulate carriers have important therapeutic application in site-speci fic drug
delivery or medical imaging. However, injected particles are generally eliminated by the reticulo-
endothelial system within minutes after administration and accumulate in the liver and spleen. To obtain a
coating that might prevent opsonization and subsequent recognition by the macrophages, sterically stabi-lized nanospheres were developed using amphiphilic diblock or multiblock copolymers. The nanospheresare composed of a hydrophilic polyethylene glycol coating and a biodegradable core in which various
drugs were encapsulated. Hydrophobic drugs, such as lidocaine, were entrapped up to 45 wt% and the release
kinetics were governed by the polymer physico-chemical characteristics. Plasma protein adsorption wasdrastically reduced on PEG-coated particles compared to non-coated ones. Relative protein amounts were
time-dependent. The nanospheres exhibited increased blood circulation times and reduced liver accumula-
tion, depending on the coating polyethylene glycol molecular weight and surface density. They could befreeze-dried and redispersed in aqueous solutions and possess good shelf stability. It may be possible to tailor
“optimal ”polymers for given therapeutic applications.
© 2012 Published by Elsevier B.V.
Contents
1. Introduction …………………………………………………….. 3 1 7
1.1. Long-circulating drug delivery systems: PEG-based coatings ……………………………… 3 1 7
1.2. Coatings obtained by polymer adsorption ……………………………………… 3 1 7
1.3. Coatings obtained by polymer chemical attachment …………………………………. 3 1 7
1.4. Mechanism of action of the PEG coating ………………………………………. 3 1 8
2. PEG-coated nanospheres ……………………………………………….. 3 1 9
2.1. Degradable PEG-Rand PEG
n-R copolymers: synthesis and characterization . . . ……………………… 3 1 9
2.2. PEG-coated nanosphere characterization ……………………………………… 3 2 0
2.2.1. Morphology ……………………………………………… 3 2 1
2.2.2. Size optimization and in fluence of the surfactant ………………………………. 3 2 1
2.2.3. Surface characterization and stability ……………………………………. 3 2 2
3. Drug encapsulation ………………………………………………….. 3 2 2
4. Interactions with blood components and macrophages …………………………………… 3 2 2
5. Biodistribution ……………………………………………………. 3 2 3
6. Perspectives and conclusions ……………………………………………… 3 2 4
Acknowledgements ……………………………………………………. 3 2 4
References ……………………………………………………….. 3 2 4Advanced Drug Delivery Reviews 64 (2012) 316 –326
☆PII of original article: 0169-409X(95)00026-7. The article was originally published in Advanced Drug Delivery Reviews 16 (1995) 215-233.
⁎Corresponding author. Tel.: +33 83175261; fax: +33 83379977.
E-mail address: Ruxandra.Gref@cep.u-psud.fr (R. Gref).
0169-409X/$ –see front matter © 2012 Published by Elsevier B.V.
http://dx.doi.org/10.1016/j.addr.2012.09.008
Contents lists available at SciVerse ScienceDirect
Advanced Drug Delivery Reviews
journal homepage: www.elsevier.com/locate/addr

1. Introduction
There has been a growing interest in the development of a colloi-
dal drug carrier which is small enough for intravenous administration
and possesses an adequate circulation half-life in order to enable drug
release into the vascular compartment in a continuous and controlled
manner.
There are numerous potential applications for such a system: the
protection of sensitive therapeutically active molecules against in
vivo degradation, the reduction of toxic side effects which can occur
when some highly active drugs like those used for cancer treatment
are administered in the form of a solution, the increase of patient
comfort by avoiding repetitive bolus injection or the use of perfusion
pumps, and the achievement of more favorable drug pharmacokinetics.
The successful delivery of macromolecular drugs is often problem-
atic[1]. Protein drugs generally cannot be administered orally, since
they are hydrolyzed or denatured in the gastrointestinal tract. Paren-
teral administration also poses some problems, since proteins are
usually quickly metabolized and eliminated. The typical pharmacoki-
netic half-lives of proteins range from 2 to 30 min [1]. The pharmaco-
kinetics of these compounds could be improved by encapsulating
them into long-circulating drug delivery devices.
The effective use of pharmacologically active substances in the
chemotherapy of cancer, viral infections, and many other diseases
suffers from non-speci fic toxicity and poor tissue speci ficity of
drugs. Polymeric nanoparticles are possible carriers for targeting
these compounds by the intravenous route in order to increase the
compound's effectiveness in the diseased tissue and reduce general
toxicity [2,3].
Various types of systems (liposomes, emulsions, micelles, and
nanoparticles) were developed over the last decade in order to
achieve controlled intravenous drug delivery or targeting to speci fic
tissues. However, most of these systems, via recognition by the phago-
cytic cells (mainly the cells of the mononuclear phagocyte system and
the polymorphonuclear leukocytes), are detected as foreign products
and quickly removed from blood circulation. Essentially, macrophages
located in the reticuloendothelial system (RES) (of which the Kupfer
cells in the liver comprise 85-95% of the total intravascular phagocytic
capacity [4]) play a crucial role to phagocytize injected particles. For
example, polystyrene (PS) particles as small as 60 nm disappear fromblood within minutes [5]. Similar short half-lives were observed
whatever the chemical composition of the injected particles: albumin
[6], poly(lactic acid) (PLA) [7,8], poly(lactic acid-co-glycolic acid)
(PLGA) [9], polycyanoacrylate [10] or polyacryl starch [11]. Therefore,
using these types of systems, it is only possible to target drugs to RES
organs, mainly the liver.
It is generally assumed that the rapid particle phagocytosis is
mediated by the adsorption of certain blood components (opsonins)
onto the surface of the particles; for example, in vitro experiments
with liposomes [12] showed that serum complement, one of the
most important components of the opsonin s-ystem of the body,
and in particular C3, strongly activates phagocytosis by macrophages.
Other complements, such as C4, also seem to be involved in the
phagocytic process [13].
The main challenge for administering particulate drug carriers
into the vascular compartment is site-selective drug delivery. Such a
carrier would avoid indiscriminate interactions with the RES, selec-
tively reach the desired tissue, release the active compound with an
optimal rate and finally degrade into non-toxic elements, which can
be eliminated from the body and which do not cause side-effects.
Various attempts were made to achieve long blood circulation times
by avoiding RES recognition, mainly by chemically attaching or adsorbing
appropriate polymers or molecules at the particle surface [5,14 –21],
which would reduce or minimize the interaction with opsonins.
After reviewing some of the long-circulating systems thus obtained
(focused on nanoparticles and liposomes), this article will discuss anewly developed type of particle with an increased blood half-life:
poly(ethylene glycol) (PEG)coated biodegradable nanospheres
[22–33].
Nanospheres (or nanoparticles) are de fined as macromolecular solid
colloidal particles of less than 1 μm, in which drugs or other biologically
active molecules are dissolved, entrapped or encapsulated, chemically
attached to the polymers or adsorbed to the particle surface [34].
1.1. Long-circulating drug delivery systems: PEG-based coatings
Particle size and shape greatly in fluence their organ distribution.
To circulate through the smallest capillaries, the particle size should
be less than 5 μm. Moreover, a diameter of less than 200 nm is
required to avoid spleen filtering effects [35]. In addition, the surface
of the particles should possess RES-avoiding properties (be stealthy
with regard to phagocytic cells). Allen [19] wrote "if you want to be
invisible, look like water"; indeed, particles with hydrophilic surfaces,
upon which water molecules can readily adsorb, have longer blood
half-lives [5,14 –20].
Particles with neutral surfaces seem the most appropriate with
regard to blood persistence [18,36] . Various attempts were directed
towards altering the carriers' surface properties to reduce their RES
clearance. This alteration was performed by adsorbing or chemically
attaching appropriate hydrophilic and neutral polymers to their surface.
1.2. Coatings obtained by polymer adsorption
In the case of PS nanoparticles, the rapid uptake by RES may be
partially overcome by coating the particles with arti ficial surfactants
such as Poloxamine 908 [5]or 1508 [14]. Poloxamer and Poloxamine
surfactants contain blocks of PEG and blocks of poly(propylene
glycol) (PPG). The more hydrophobic PPG blocks adsorb on the PS
surface and the more hydrophilic PEG blocks stick out of the surface
and form a protective coating [37]. PEG was adsorbed into carboxylated
PS and poly(styrene-butadiene)latexes from water [38].
The adsorption was kinetically controlled and decreased with
an increase in surface roughness and polarity. However, non-
degradable PS particles are not realistic therapeutic systems. Similarly
to PS particles, poly(methyl-methacrylate) colloidal carriers coated
by Poloxamer [39] circulate longer in blood.
The polymer adsorption approach was adapted from model PS
particles to degradable PLGA particles [40] coated by adsorption of
amphiphilic diblock PEG-PLA diblock copolymers, which advanta-
geously replaced the non-degradable surfactants previously used
(Poloxamer and Poloxamine) [5,14] . However, a comparison of the
in vivo results obtained with the two systems, PLGA/PEG-PLA and
PS/Poloxamer 908, suggested that the adsorbed layers of PEG-PLA
might have a lower stability or resistance to protein adsorption than
Poloxamer 908 ones.
The adsorbed polymers might desorb in vivo, due to replacement
by blood compounds with a higher af finity for the particle surface
[18], and this would pose problems for their intravenous administra-
tion. To increase the coating stability, polymer chemical attachment is
an alternative to polymer adsorption.
1.3. Coatings obtained by polymer chemical attachment
Model spherical PS particles with covalently bound PEG2K
(molecular weight 2000 Da) chains on their surface were less seques-
tered by the liver than PS ones [41]. Moreover, a correlation was
found between the PEG surface density and the blood half-life. At the
highest PEG surface density, the increase in blood circulation time was
higher than that obtained with particles coated by Poloxamine 908
adsorption [41].
PEG 350 was attached to the surface of PS latex particles [42].B y
assuming that all the PEG present in the system was located at the
surface, it was calculated that one PEG molecule covers about 1 nm2.317 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

PEG was covalently attached to the surface of poly(alkyl cyanoac-
rylate) particles [43] or to the surface of chemically cross-linked
albumin nanospheres [20]. In the latter case the particles' uptake
by cell culture macrophages was signi ficantly reduced. Poly(aspartic
acid) polymers with pendant PEG chains and adriamycin (a hydro-
phobic drug) chemically attached to them formed long-circulating
micelles [44]. Covalently bound to proteins, PEG forms macromolecu-
lar conjugates with longer blood circulation time than native proteins
and with reduced immunogenicity and antigenicity [45]. Attached to
surfaces, PEG was shown to increase their biocompatibility [46] and
reduce the adsorption of proteins and blood components [47–51].
When appropriate polymers are attached to their surface, liposomes
are an important class of long-circulating particles, because they are
generally composed of substances naturally occurring in the body.
Although liposomes were hypothesized to make good drug delivery
systems more than 20 years ago [52], in some cases they have circula-
tion half-lives as short as a few minutes, as a result of their rapid uptake
into the cells of the RES. Thus, many of their therapeutic applications
have been limited to the delivery of drugs to this system; liposomes
were used to treat leishmaniasis [53] or to manipulate macrophage
function, for example by delivering immune modulators [54],a m o n g
other applications. To enable additional applications, such as the deliv-
ery of drugs to tumor cells, liposomes with increased blood half-lives
were designed (termed "Stealth" or sterically stabilized), by attempting
to mimic the surface structure of blood cells. For example, carbohy-
drates [55] and polysaccharides such as monosialoganglioside GM1
[56] were attached to the liposomes' surface and a signi ficant increase
in blood circulation times was achieved. A coating with PEG not only
dramatically increased the liposomes' blood half-life [15,16] ,b u ta l s o
solved some problems related to the use of GM1, such as its costliness
and its dif ficulty in puri fication. So far, a PEG coating appears to be supe-
rior to other polymers in conveying signi ficant advantages to liposomes.
A few of these advantages are: decreased RES uptake, increased blood
residence time, increased stability to contents leakage, increased flexi-
bility in lipid composition, and dose-independent pharmacokinetics
[19].
However, the amount of PEG which can be included in the
liposome lipid bilayers decreases with increasing PEG molecular weight
[57]; from 15 mol% in the case of PEG 120 to 5 –7 mol% in the case of
PEG2K and PEG5K. Above 7.5 mol% of PEG 1900, a liposome dissolutionwas observed [58].
The bilayer rigidity is an essential factor to prolong the liposome
circulatory half-life, for example by decreasing the interactions with
plasma proteins. A low bilayer rigidity can lead to destabilization of
the liposome membrane and release of encapsulated drug, thus
aborting the function of the microparticle delivery system [59,60] .
1.4. Mechanism of action of the PEG coating
The research work summarized in the previous paragraphs shows
that hydrophilic PEG-based coatings signi ficantly increase the nano-
particle or liposome blood circulation time. Among hydrophilic
polymers, PEG has the advantage to be considered non-toxic and
was approved by the Food and Drug Administration (FDA) for inter-
nal use in humans [61]. The blood half-life of PEG extended from
18 min to 1 day as the PEG molecular weight increased from 6 K to
190 K [62].
Unraveling the mechanistic action of hydrophilic and in particular
PEG-based coatings would be a step forward in the design of particles
with optimized surface properties. A vast amount of research work
has been devoted to a better understanding of the mechanism of
the PEG coating utilized to extend particle blood circulation times. A
generally assumed mechanism is based upon the formation of a
sterically hindered, hydrophilic coating which avoids opsonization
by plasma proteins [63–65]. The hydrophilicity was considered a
main requirement, but it turned out not to be a suf ficient one. Indeed,liposomes were coated with series of hydrophilic polymers, among
which maltopentaose, estimated to be more hydrophilic than PEG5K,
but they were still removed from blood circulation very rapidly in
mice [65]. Dextran-coated liposomes circulate shorter than PEG-
coated ones [66], in spite of the more hydrophilic nature of dextran
compared to PEG.
It has been proposed [63,65] that besides hydrophilicity, chain
flexibility is another major feature necessary for the coating polymers
to provide long-circulating particles. Due to the transient, flexible and
rapidly changing structure of PEG, the immune system would have
difficulties in modeling an antibody around it [67]. The protective
layer of PEG is considered as a “cloud ”of possible chain conforma-
tions, with a density high enough to prevent the interactions of opso-
nins with the surface of the particles; only if a polymer chain posesses
both hydrophilicity and flexibility properties (to enable a high num-
ber of possible chain conformations), can it serve as an effective
coating protector for particles against opsonization [63]. Besides PEG,
other possible candidates were considered, such as poly(acrylamide)
or poly(vinylpyrrolidone). Preferably, the protective polymers should
not contain hydroxyl groups (like polysaccharides), which are targets
for complement C3, or amine groups (like polylysine), which are targets
for C4 [63].
The conditions that lead to protein repulsion from hydrophobic
plane surfaces to which PEG chains were attached to one chain end
in a "brush" con figuration were recently studied [68,69] . These
authors elaborated a mathematical model taking into account the
four types of interactions between a protein and a hydrophobic sub-
strate ( Fig. 1 ). The best conditions for protein repulsion were found
to be long PEG chain length and high surface density [68]. Let D
be the distance between the anchorage to the substrate of two termi-
nally attached PEG chains (cf. Fig. 1 ). In the case of small proteins
(approximated as spheres with a radius of 2 nm), Dshould be around
1 nm, whereas for larger proteins (6 –8 nm), Dshould be around
1.5 nm.
It has been suggested that both the reduction in the adsorption
of opsonins and the selective adsorption of certain plasma components
(dysopsonins) prevent the recognition and uptake of nanoparticles
by macrophages; the competition between these two mechanisms is
believed to be the key in controlling the particle uptake by macrophages,
and hence their biodistribution [70].M u i re ta l . [71]suggested that two
serum components (one with a molecular weight below 30 000 Da
and the other with a molecular weight higher than 100 000 Da) are
the principal factors which result in a dysopsonic action.
The interactions between injected particles and blood compo-
nents are complex. These compounds might reversibly or irreversibly
adsorb on the surface of the particles, and might be replaced by
Fig. 1. Interactions between a protein and a hydrophobic substrate with attached PEG
chains (adapted from [68]). 1: hydrophobic attraction between the protein and the
substrate; 2: steric repulsion resulting from PEG chain constriction; 3: van der Waalsattraction between the protein and the substrate; 4: van der Waals attraction betweenthe protein and the PEG chains.318 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

others. Two-dimensional polyacrylamide gel electrophoresis (2-D
PAGE) will probably be a helpful tool to gain insight into these
phenomena [72].
2. PEG-coated nanospheres
Even if the exact mechanism which leads to an increase in the
blood half-life of PEG-coated surfaces has not been entirely elucidated,
the sum of the research work presented in the previous paragraphs
clearly shows that a PEGbased coating (especially if obtained by chem-
ical attachment) is a key in the design of long-circulating particles, and
it has inspired recent approaches to design PEG-coated nanoparticles
[22–33].
As model PS particles, these nanospheres have compact polymeric
cores (in order to ensure good stability) but the cores are degradable.
Composed of poly(hydroxy acids) such as PLA or PLGA, the cores
should degrade by hydrolysis into harmless elements which could
be excreted (such as lactic and glycolic acids), in order to avoid parti-
cle accumulation in the body. Investigations of the potential use
of PLA for medical purposes date from 1966 [73]. PLA, PLGA and
polycaprolactone (PCL) showed a good ability to encapsulate various
drugs. By adjusting the physico-chemical characteristics of these
polymers, different encapsulation properties and release kinetics
could be obtained [3].
As with PEG-coated liposomes, the PEGcoated nanospheres have
PEG chains attached to their surface at one chain end, in a brush con-
figuration (as in Fig. 1 ), which should avoid or reduce the interactions
with blood proteins and therefore impart RES-avoiding properties. To
achieve the core-shell structure described above, block amphiphilic
polymers of the type PEG-R were synthesized. “R”is chosen among
the bioerodible polymers cited above (PLA, PLGA, PCL). The two
blocks have a tendency to easily phase-separate in the presence
of water [74] and have different solubilities in water and organic
solvents. This property was used to obtain the core-shell structure
by an o/w emulsi fication procedure ( Fig. 2 ). For this, PEG-R polymers
were dissolved in an organic solvent immiscible with water (such as
ethyl acetate or methylene chloride). The o/w emulsion was formedin an aqueous phase, and the organic solvent was allowed to slowly
evaporate. This led to a progressive increase in polymer concentration
inside the droplets. R is insoluble in water, but highly soluble in the
organic solvent; conversely, PEG is highly water-soluble, soluble in
methylene chloride, and practically insoluble in ethyl acetate. This
leads to a tendency of PEG chains to migrate towards the water
phase to form sterically stabilized particles ( Fig. 2 ), with a core presum-
ably mostly composed of R chains. After complete solvent evaporation,
t h en a n o s p h e r ec o r es o l i d i fies, thus entrapping the hydrophobic biolog-
ically active molecules.
2.1. Degradable PEG-Rand PEG
n-R copolymers: synthesis and
characterization
One research trend is to synthesize new tailored biocompatible
polymers. For example, triblock R-PEG-R polymers were synthesized
to take both advantage of the softness and flexibility of the PEG blocks
and of the rigidity and hydrophobicity of the “R”blocks (PLA or PCL),
in order to obtain new materials with desired mechanical and degra-
dation behavior [75–82].
To prepare PEG-coated nanospheres, PEG-R polymers were syn-
thesized by ring opening polymerization of monomers (lactide and/
or glycolide, caprolactone) in the presence of monomethoxy PEG
(MPEG) [22,24,30] . The catalyst chosen was stannous octoate, widely
used to prepare PLA polymers and approved by the FDA as a food
stabilizer [83]. Toxicological data are available for this catalyst [84].
Moreover, among different catalysts used, stannous octoate was the
most effective to synthesize PLA-PEGPLA polymers [78].The polymerization reaction was followed by gel permeation
chromatography to determine molecular weight and polydispersity
[25].Fig. 3 shows typical gel permeation chromatograms of reaction
mixtures after various times (starting material: MPEG, lactide and
glycolide). The consumption of monomers (lactide and glycolide)
is observed from a decrease of peak “D”. Simultaneously, peak “P”
(corresponding initially to MPEG) shifts towards lower retention
times (higher molecular weights), indicating that an addition reac-
tion takes place at the hydroxyl end group of MPEG. Finally, only
one peak is obtained, corresponding to diblock PEG-R polymers.
After puri fication, the exact chemical composition of the polymers
was determined by1H- and13C-NMR spectroscopy [22].
The crystallinity and glass transition temperature ( Tg) of a series
of diblock PEG-PLGA polymers were determined by differential scan-
ning calorimetry (DSC) [25] (Fig. 4 ). The first heating trace shows an
endothermic peak corresponding to the fusion of the crystallites in
the sample; the melting enthalpy corresponds to the total amount
of PEG in the sample, indicating that a phase separation occurred
between PEG and PLGA. An X-ray study conducted with these samples
Fig. 2. Schematic representation of the nanosphere fabrication procedure following an
emulsion-solvent evaporation procedure.
Fig. 3. Time course of the polymerization reaction between MPEG and lactide/glycolide
followed by gel permeation chromatography (adapted from [25]). Peak P: PEG-PLGA
copolymer; peak D: starting monomers (lactide/glycolide).319 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

confirmed that only PEG crystallizes in the samples; the diffraction
chromatograms showed only speci fic peaks corresponding to PEG
crystallites, the PLGA polymers being amorphous [22].
After the first run, the samples were quenched by rapid cooling
and presented an amorphous structure. The second heating trace
at 10 °C/min did not show a melting endotherm but a single Tg
(Fig. 4 ). We suppose that the single Tgobserved in this case is due
to an entanglement effect of PEG and PLGA chains, that kinetically
could not phase-separate during the rapid cooling. The calculated Tg
follows the Fox law for polymer blends:
1=TgðȚPEG/C0PLGA¼w=TgðȚPEGțw=TgðȚPLGA
where wis the weight fraction of each polymer block.
Multiblock PEG n-Rmcopolymers were also synthesized [26,29] .
First, several monomethoxy monoamine PEG (MPEG-NH 2) chains
were attached together at one chain end by reaction with citric,
mucic, tartaric acid or other polyfunctional molecules. The remaining
hydroxyl groups were further used to initiate the ring-opening poly-
merization of “R”. The structure of the polymers thus obtained (when
R=PLA) is indicated in Scheme 1 .
The morphology, degradation and drug encapsulation behavior
of copolymers containing PEG blocks strongly depends on their
composition. For example, poly(L-lactic acid)-PEG-poly(L-lactic acid)
(PLLA-PEG-PLLA) polymers crystallize forming spherulites with a mor-
phology depending on copolymer composition, hydrolysis time and
crystallization temperature [85]. The degradation of poly(D,L-lactic
acid)-PEG-poly(D,L-lactic acid) triblock copolymers was studied [74].
The rate of water uptake in films prepared of these polymers was
biphasic, suggesting two possible mechanisms. Compared to PLA,
PLA-PEG-PLA copolymers showed faster degradation kinetics, but the
overall biological response to both types of implants was good and com-
parable in both cases [86].
The diffusion of dyes through degraded semicrystalline films
prepared from PLLA-PEG-PLLA polymers was further studied [87].
The diffusion rate increases as the PEG content in the copolymers
increases. The steady state of mass flux was not reached over a diffu-
sion time of 1000 h, because of polymer degradation.
Bovine serum albumin was encapsulated in PEG-PLA and PEG-PLGA
[88] and PLA-PEG-PLA or PLGA-PEG-PLGA [89] microspheres. Striking
differences were observed with the respective homo- or copolymers
with regard to microsphere morphology, degradation behavior, and
drug release patterns.
To form PEG-coated nanospheres, diblock PEG-R polymers were
used and not triblock R-PEG-R polymers. R-PEG-R would lead to the
formation of PEG coils on the surface with two anchoring points to
the core. For the same PEG chain length, the maximum coating thus
obtained would be half as thick as the one obtained when diblock
polymers are used. As a result, the PEG coating on R-PEG-R
nanospheres would be less effective than the one on PEG-R particles,
in order to prevent opsonization and increase blood circulation times.
2.2. PEG-coated nanosphere characterization
The PEG-R nanospheres were prepared according to the procedure
described in Section 2 and schematized in Fig. 2 . Hydrodynamic
diameter and size distribution, some of the most prominent features
of nanoparticles, were determined by quasi-elastic light scattering,
(QELS) [22,23] and compared to those obtained from morphology
studies (essentially using electron microscopy). The nanospheres
Fig. 4. DSC thermograms of PEG5K-PLGA copolymers with increased chain length of
PLGA. The first run (heating rate 10°C/min) was obtained with puri fied polymers.
The samples were rapidly quenched and a second run (10°C/min) was enregistered.
Scheme 1. Chemical structure of some diblock PEG-R and multiblock PEG n-Rm
copolymers used for the preparation of PEG-coated nanospheres.320 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

recovered by centrifugation, after preparation, can be lyophilized and
redispersed in aqueous solutions and they presented similar size dis-
tributions according to QELS studies [22]. Nanosphere powders show
good shelf storage properties.
2.2.1. Morphology
The spherical shape of PEG-coated nanospheres was validated by
scanning electron microscopy (SEM) ( Fig. 5 ) and atomic force micros-
copy (AFM) [22], a non-destructive technique. SEM needs a previous
coating of the sample with gold and/or palladium, whereas AFM allows
a higher resolution without the need of coating the nanospheres.
In addition to these techniques, freeze-fracture electron microscopy
was used [23]to allow observation of the nanospheres surface ultrastruc-
ture ( Fig. 6 ). The samples were frozen, fractured, shadowed by platinum
and carbon and the replicas were observed by transmission electron
microscopy. Fig. 6 shows typical images of PLGA, PEG20K-PLGA, and
lidocaine-loaded PEG20K-PLGA nanospheres. Hemispheric replicas of
nanospheres were found in all samples. PLGA nanospheres had a smooth
surface, whereas both unloaded and lidocaine-loaded PEG20K-PLGA
nanospheres showed typical particles on their surface. These particles
with an average greatest diameter of about 12 nm were attributed to
the presence of PEG chains at the surface of nanospheres.
PEG20K-coated nanospheres containing 50 wt% lidocaine, a crys-
talline drug, consistently showed filamentous structures, in addition
to the surface particles also found in samples without lidocaine
(Fig. 6 ). These filaments of about 11 nm apparent diameter could be
attributed to small lidocaine crystals inside the nanospheres. This
hypothesis is in agreement with previous DSC studies performed on
nanospheres, which showed lidocaine melting endotherms in these
samples [22].
2.2.2. Size optimization and in fluence of the surfactant
The nanosphere hydrodynamic diameter dand size distribution were
measured using quasi-elastic light scattering (QELS). This technique was
used to follow the nanosphere aggregation and de-aggregation behav-iors in aqueous solutions [22].
The in fluence of different fabrication parameters on dwas studied
with the aim to obtain particles of less than 200 nm [28]. The size
of the PEG5-PLA45 nanospheres was reduced by increasing theconcentration of the surfactant used for their preparation, poly(vinyl
alcohol) (PVA) or cholic acid, sodium salt (CA). ddecreases with
reducing the polymer concentration in the organic phase. These
results are in agreement with those reported in size optimization
studies on PLGA nanospheres [90].
Several surfactants were tested with regard to their ability to form
nanospheres without retaining their composition [28]. Among the
anionic biological surfactants tested, CA was the most effective to
reduce both hydrodynamic diameter and polydispersity index ( Table 1 ).
Only a low amount (less than 6 wt%) of this compound remained associ-
ated with the nanospheres after their preparation. In contrast, when PVA
was used as a surfactant, a very high amount of it remained in the
nanosphere composition, possibly by entanglement with PLA chains in
the core. This hypothesis was supported by the fact that, after repetitive
washing steps in distilled water, the residual PVA amount still repre-
sented as high as 30-50% of PLA or PEG-PLA nanospheres weight.
Some of the surfactants tested were unsuccessful for the prepara-
tion of PLA nanospheres, whereas they could be used to prepare
PEG-PLA ones. This result was attributed to an additional steric stabi-
lizing effect of PEG chains on the surface of the latter particles.
Whatever the surfactant used, PLA nanospheres have a negative
surface charge, as indicated by the negative values of the zeta poten-
tial (ZP) ( Table 1
), ranging from −4 mV (when the non-ionic PVA
surfactant was used) to −12 mV (when anionic surfactants were
employed). The negative surface charge was attributed to the pres-
ence of carboxyl end groups (from PLA), located near the surface
Fig. 5. Scanning electron microscopy of PLGA40K (A) and PEG5K-PLGA45K (B) nanospheres.
Fig. 6. Freeze-fracture electron microscopy of PLGA40K (A), PEG20K-PLGA180K (B)
and lidocaine-loaded (45 wt%) PEG20K-PLGA180K (C) nanospheres.321 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

and to the presence of adsorbed residual surfactants. In contrast, the
ZP of PEG-coated nanospheres was practically zero, regardless of the
surfactant used. This result was attributed to a complete screening
of the surface charge by the PEG chains covalently attached to it.
The PEG coating layer shifts the shear plane of the diffuse layer to a
larger distance from the nanospheres, and this results in a decrease
of the measured absolute value of ZP.
The ZP studies are in agreement with previous ones dealing with
PS[21] or poly( β-malic acidco-benzyl malate) [91] particles coated
by Poloxamer or Poloxamine adsorption. The adsorption of these
polymers containing PEG blocks resulted in an increase in the surface
hydrophilicity and a reduction of the ZP. Similarly, these two charac-
teristics were modi fied by the covalent attachment of PEG to the
surface of albumin nanoparticles [20], with an initial ZP of −37 mV.
A coating with MPEG 750 did not modify ZP, but a coating with
MPEG2K increased ZP up to −33.5 mV and a coating with MPEG5K
up to−23 mV. This gradual modi fication of ZP was attributed to an
increase in the thickness of the PEG coating layer with the chain length
of PEG. After a complete screening of the surface charge by additional
PEG attachment, the resulting ZP was −5 mV, which was considered
sufficient to avoid RES recognition [20].
2.2.3. Surface characterization and stability
X-Ray photoelectron spectroscopy (XPS) was used to determine
the relative amount of PEG in the first 5 –10 nm surface layers of
PEG on the surface of PEG-R and PEG 3-PLA nanospheres and to follow
its stability during degradation at 37°C in water [22–24,26] . The coat-
ing was stable during a 24 h period, in agreement with the dosage of
the amount of PEG detached in the supernatant [24]. For example,
less than 8% of PEG was detached from PEG5K-PLGA45K nanospheres
in the first 60 h of incubation in 0.1 M phosphate buffer solutions
(PBS) at 37°C ( Fig. 7 ). In the case of PEG12K-PLGA100K nanospheres,
less than 10% PEG was detached in the first 60 h of incubation under
the same conditions. Over 3 weeks, the nanosphere mean diameter
(as determined by QELS) remains pratically constant, then drops
down suddenly, presumably when water-soluble oligomers are
formed [23].
Intensive studies are underway to develop an exact characteriza-
tion of the surface of PEG-coated nanoparticles, in particular, to ob-
tain the apparent hydrodynamic layer thickness values. PEG-based
coatings modify particle surface composition, charge, and hydropho-
bicity. Review articles [3,21] summarize some of the most widely
used techniques for speci fic tasks: to determine surface chemical
composition (mainly X-ray photoelectron spectroscopy and second-
ary ion mass spectroscopy (SIMS)), hydrophobicity (hydrophobic in-
teraction chromatography; Rose Bengal binding method; partitioning
between two aqueous phases dextran/PEG; contact angle measure-
ment), and surface charge (electrophoresis; laser Doppler anemome-
try; amplitude weighted phase structuration).3. Drug encapsulation
Various drugs were successfully entrapped inside PEG-coated
nanospheres: lidocaine [22,26] , prednisolone, carmustine [26],i b u p r o f e n
[30], and oligonucleotides [33]. To entrap these active compounds, the
incorporation method schematized in Fig. 2 was used; drug and polymer
were dissolved together in an organic solvent, and then an o/w emulsionwas formed by sonication or micro fluidization.
Lidocaine was used as a model hydrophobic drug in an attempt
to establish the main factors which determine the encapsulation
efficiencies, nanosphere mean diameter and release patterns [22,26] .
For these studies, drug-loaded nanospheres were prepared from
various diblock PEG-R and multiblock PEG
n-R copolymers, where R
was chosen among PLA, PLGA, PCL or PSA (the structure of these poly-
mers is given in Scheme 1 ). Regardless of the polymer used, high load-
ings and entrapment ef ficiencies were obtained. The nanosphere mean
hydrodynamic diameter and the entrapment ef ficiency strongly depend
upon the polymer physico-chemical characteristics. The hydrodynamic
diameter increases with the PEG and R molecular weights and is practi-
cally independent of the drug loading. In the case of diblock PEG-PLGA
polymers, the entrapment ef ficiency is practically independent of the
PEG molecular weight. Conversely, in the case of PEG 3-PLA polymers,
a slight decrease of the entrapment ef ficiency with the PEG molecular
weight was observed [26].
Drug release kinetics were also in fluenced by the nanosphere drug
loading ( Fig. 8 ). The possibility of lidocaine crystallization occurring
inside the nanospheres at high loadings (above 30 wt%) has been
suggested [22]. Part of the drug would homogeneously mix with the
chains in the core, and part of it would coexist as crystallites.
Freeze-fracture electron microscopy of nanospheres at high lidocaine
loadings ( Fig. 6 ) tend to support this hypothesis.
14C-Ibuprofen, a drug with a half-life of only a few minutes, was
entrapped in PLGA and PEG2K-PLA30K nanospheres. The plasma
half-life was increased in the case of PEG2K-PLA30K particles up to
2.5 h [30].
Prior to their encapsulation into PEG-PLA nanospheres, oligonu-
cleotides were associated by ion pair formation to oligopeptides
[33]. The resulting globally neutral complex could be entrapped
with high ef ficiencies into the nanospheres.
4. Interactions with blood components and macrophages
The analysis of plasma protein adsorption on nanoparticles has been
established using two-dimensional polyacrylamide gel electrophoresis
(2-D PAGE) [72]. The nanoparticles are incubated in vitro in human plas-
ma and then separated from the plasma by centrifugation. The adsorbed
proteins are washed off the particles and analyzed electrophoretically.Table 1
Influence of the surfactant on the hydrodynamic diameter (d), polydispersity index
(PI) and zeta potential (ZP) of PEG5KPLA45K and PLA100K nanospheres.
Surfactant PEG5-PLA45 PLA100
cmc
(mM)d
(nm)PI ZP
(mV)d
(nm)PI ZP
(mV)
Cholic acid, Na salt 14 148 0.049 1.1 117 0.021 -9.3
Glycocholic acid, Na salt 7.1 154 0.084 0.7 130 0.082 -12.3Taurocholic acid, Na salt 3-11 164 0.078 0.3 135 0.067 -12.4
Deoxycholic acid, Na salt 5 343 0.6 ND >800 – –
Poly(vinyl alcohol) 192 0.138 -1.2 203 0.134 -4.2
88% hydrolyzed, 13-23 K
n-Octyl- β-
D-glucopyranoside 20-25 232 0.226 ND – – –
Surfactants were used at the critical micellar concentration (cmc), with the exception
of PVA (used at 0.6 wt% in water). ZP was measured with washed nanospheresresuspended in sodium chloride 10
-3M.
Fig. 7. Fraction of PEG detached from PEG5K-PLGA45K nanospheres during incubation
at 37°C in phosphate buffer solutions (pH 7.4).322 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

Recent studies with an incubation time of 5 min showed that on
PLGA45K nanospheres major proteins were the apolipoproteins J,
C-III, E and A-I ( Fig. 9 ). Additionally, for example, considerable
amounts of fibrinogen, immunoglobulin G (not shown) and apolipo-
protein A-IV were detected. The total amount of adsorbed proteins
on the PLGA particles was more than 4 times higher than on all
PEG-PLGA nanospheres which were investigated ( Fig. 9 ). The total
protein amount was similar on PEG5K-PLGA45K, PEG12K-PLGA100K
and PEG20K-PLGA180K.
Especially the amounts of apolipoprotein J and C-III were drasti-
cally reduced on all PEG-PLGA particles ( Fig. 9 ). The adsorption of
fibrinogen and immunoglobulin G was also decreased by the PEG
content in the particles. The latter effect is well in agreement with
observations of the protein adsorption on PS particles which were
surface-modi fied with poloxamers [92].
The different PEG-PLGA particles showed similar protein adsorp-
tion patterns. Nevertheless, some differences could be detected. For
example, PEG5K-PLGA45K particles showed a higher adsorption of
apoA-IV than the other PEGPLGA particles. On PEG12K-PLGA100K
the highest amount of apoE was detected and on PEG20K-PLGA
180 K the amount of apoC-III was slightly higher than on the other
PEG-PLGA particles.
Prolonging the incubation time of the particles in the plasma led to
considerable changes in the adsorption patterns. After an incubation
time of 20 h, apolipoprotein C-III was strongly reduced; it was barely
detectable on any of the investigated particles. In contrast to that, the
amounts of apoA-IV and apoE were distinctly increased on all the
particles, especially on the PLGA nanospheres. The time dependence
of adsorption of these proteins was very similar to that on PS particles
coated with poloxamers [92].
In general, a distinct reduction of protein adsorption was achieved
by employing PEG-PLGA polymers for particle production as com-
pared to pure PLGA, which is in agreement with the generally accept-
ed low protein adsorption of PEG surfaces [93]. The qualitative
composition of the adsorbed proteins was changed as well by the in-
troduction of PEG. The differences in the adsorption patterns of the
particles with varying PEG contents were probably caused by differ-
ent PEG surface density rather than different PEG chain length
[68,92] .
The relevance of the adsorbed plasma proteins to the in vivo fate
of the particles still remains to be uncovered. It has to be kept in
mind that not only single proteins but also the composition, i.e., the
ratios of the adsorbed proteins and their conformation may be of im-
portance. Special interest has to be directed to the time dependence
of adsorption.
PLA70K nanospheres strongly activate the complement system,
whereas PEG2K-PLA30K ones hardly do [94]. PLA and PEG2K-PLA30K
nanoparticles of identical size (160 nm) were incubated with human
platelet-poor plasma [32]. In the case of PLA particles, a diminution of
activity of factor V was observed, together with particle aggregation.Conversely, all tests performed on PEG2K-PLA30K particles (concerning
global coagulation times and the activity of different plasma coagulation
factors) were identical to control samples [32]. The observed absence of
PEG2K-PLA30K particle aggregation in plasma is a necessary condition
for their intravenous administration.
PEG2K-PLA and PEG5K-PLA nanospheres showed an ability to avoid
in vitro macrophage recognition [31]. PEG5K-PLA30K particles were
phagocytized 9 times less than PLA ones. PEG5K was more effective in
preventing macrophage uptake than PEG2K.
All these results suggest the value of PEG-PLA nanospheres, as op-
posed to PLA nanospheres, for use as injectable carriers able to escape
macrophage recognition. In further studies, the establishment of the
correlation of surface properties of the particles with protein adsorp-
tion and subsequent in vivo behavior will be approached.
5. Biodistribution
PEG-coated and non-coated nanospheres were injected into BALB/c
mice [22] to compare their biodistribution. The particles were core-
labeled with111In, a radioactive γ-emitting compound. For this,
diethylene triamine pentaacetic stearyl amide (DTPA-SA) was labeled
with111In by the transchelation method [95] and encapsulated into
the particles during their formation. Due to the hydrophobic character
of DTPA-SA, the label was firmly attached to the nanoparticles. It was
possible to observe images by γ-scintigraphy of the mouse body after
injection of PLGA and PEG20K-PLGA nanospheres [22]. Fifteen minutes
after injection, PLGA particles accumulate in the liver and spleen,
whereas PEG20K-PLGA nanospheres circulate well and radioactivity in
the blood pool (heart and lung) was much higher.
In another series of experiments, nanospheres were injected in
mice which were sacri ficed and blood and organ samples were counted.
Five minutes after injection of uncoated PLGA nanospheres, 66% of
nanosphere associated111In radioactivity was found in the liver and
about 5% in the blood. The results were reversed in the case of
PEG20K-coated particles: 17% of injected radioactivity was in the liver
and more than 40% in the blood [22]. As expected, non-PEG protected
PLGA nanoparticles immediately accumulate in the liver; conversely,
5 h after injection, PEG20K-PLGA nanospheres in the liver did not
exceed 30%. In the case of PEG-coated nanospheres, blood circulation
times increased with the molecular weight of the protecting PEG chains.
In addition, the mean diameter and the size distribution have a
significant effect on biodistribution; larger particles can be removed
by a filtering effect, whereas small particles can be subject to
Fig. 8. Lidocaine release from PEG20K-PLGA180K nanospheres (10 and 33 wt% loading)
(reproduced with permission from [22]).
Fig. 9. Plasma protein adsorption on PLGA40K (A) and PEG5K-PLGA45K (B) nanospheres
(sample preparation and 2-D PAGE protocol after [72]). Close-up of the bottom left part
of the 2-D PAGE gels. The proteines are separated on the basis of their molecularweights (MW) and isoelectric points (p I). (1) ApoA-IV, (2) ApoJ, (3) ApoE, (4) ApoA-1,
(5) ApoC-III, (6) ApoC-II, (7) ApoA-II.323 R. Gref et al. / Advanced Drug Delivery Reviews 64 (2012) 316 –326

extravasation. For example, in spite of their hydrophilic coating,
model PS particles modi fied with surface adsorbed block copolymers
were largely captured by filtration in the red pulp of the spleen when
their size was larger than about 250 –300 nm [35]. A spleen filtering
effect was also observed with PEG-PLGA and PEG 3-PLA particles in
this size range [96].
6. Perspectives and conclusions
PEG-coated nanospheres are a newly developed drug delivery
system which has potential applications for intravenous drug admin-
istration. Intensive studies are underway to optimize and explore
the possibilities offered by this recent system. For example, the
biodistribution and the drug encapsulation properties of nanospheres
prepared from new diblock and multiblock copolymers are being
studied.
The PEG coating provides protection against interaction with blood
components and thus prolongs blood circulation time by reducing
the particle capture by the RES cells. The degradable core of the
nanospheres was formed using a variety of polymers, such as PLA,
PLGA, PCL or PSA. Various drugs were encapsulated in the core and
released continuously. For a given drug, the release kinetics and theentrapment ef ficiency were governed by the polymer physico-
chemical characteristics, such as chemical structure, molecular weight,
and crystallinity. It may therefore be possible to tailor an “optimal ”
polymer for different therapeutical applications.
Long-circulating PEG-coated nanospheres could function as circu-
lating depots. By slowly releasing the active compound in the blood,
the drug plasma concentration pro files could be altered, which results
in therapeutic bene fits, such as a decreased systemic toxicity. Circulating
particles acting as depots might present advantages for the administra-
tion of drugs, such as anticoagulants, which might cause bleeding
when entrapped in local depots (such as intramuscular inserts). The
biodistribution of anti-cancer drugs (e.g., adriamycin) might also be
altered, potentially minimizing problems of cardiotoxicity.
Similarly, a possible use of PEG-coated nanospheres could be
the encapsulation of various contrast agents. A number of contrast-
enhancing agents have been developed to resolve and contrast tissues
for diagnostic imaging (e.g., γ-camera, nuclear magnetic resonance
imaging, computed tomography). Gadolinium diethylenetriamine
pentaacetic acid encapsulated in liposomes improved the detection
of liver metastases in vivo, compared to the free contrast agent [97].
Similarly, PEG-coated nanospheres could be good candidates to
encapsulate contrast agents with short blood half-lives in order to
alter their biodistribution.
PEG-coated nanospheres represent interesting alternatives as drug
delivery systems to PEG-coated liposomes. Their similar size enables
similar uses, but the nanospheres may have longer shelf lives and
avoid destabilization by interaction with serum components. Moreover,
many coating materials such as surfactants might lead to the disintegra-
tion of liposomes, but do not affect the nanospheres' integrity. The
nanospheres can also be freeze-dried and reconstituted in aqueous
solutions.
Another advantage that could be taken from the stability of
PEG-coated nanospheres is the possibility of attaching antibodies or afragment of them to the surface of the particles, without destabilizing
them, in order to achieve site-speci fic drug delivery, a major challenge
for drug administration. Ideally, these "magic missiles" [98]would accu-
mulate at the diseased tissue and locally liberate the necessary amount
of drug.
Acknowledgements
We acknowledge the fruitful collaboration and discussions with
Dr. V. Torchilin and Dr. V. Trubetskoy from Massachusetts General
Hospital, Charlestown, MA, USA, where the in vivo studies on PEG-coated nanospheres were carried on. We thank our collaborators,
Dr. Y. Minamitake (Suntory Limited, Japan) and M.T. Peracchia
(Parma University, Italy), who provided some of the key contribu-
tions reviewed here. Dr. R. Gref wishes to acknowledge the French
Foreign Affairs Ministry for a Lavoisier grant, and Dr. R. Langer
acknowledges NIH grant GM26698.
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