Polymeric nanocapsules ultra stable in complex biological media

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Polymeric nanocapsules ultra stable in complex biological media C. Rodriguez-Emmenegger a,b,∗ , A. Jäger a , E. Jäger a , P. Stepanek a , A. Bologna Alles b , S.S. Guterres c , A.R. Pohlmann c,d , E. Brynda a a

Institute of Macromolecular Chemistry, Academy of Sciences of the Czech Republic, v.v.i., Heyrovsky Sq. 2, 162 06, Prague, Czech Republic College of Engineering, Universidad de la Republica, Julio Herrera y Reissig 565, PS 11300 Montevideo, Uruguay c Faculdade de Farmácia, Universidade Federal do Rio Grande do Sul, Av. Ipiranga, 2752, BR 90610-000, Porto Alegre, Brazil d Departamento de Química Organica, Instituto de Química, Universidade Federal do Rio Grande do Sul, Porto Alegre, Brazil b

a r t i c l e

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Article history: Received 24 August 2010 Received in revised form 4 December 2010 Accepted 7 December 2010 Available online xxx Keywords: Core/shell nanocapsules Biodegradable Antifouling coating Surface initiated-ATRP Interaction with proteins

a b s t r a c t Non-specific protein adsorption from complex biological media, especially from blood plasma, is an urgent challenge for the application of nanoparticles as delivery systems, diagnostics, and other biomedical application. Nanocapsules (NC) prepared from FDA-approved degradable poly(␧-caprolactone) shell and Mygliol 812® oil in the core were coated with mono-methoxy terminated oligo(ethylene glycol) methacrylate (poly(MeOEGMA)) polymer brush layers with a well-controlled thickness at the nanometer scale up to 350 nm using surface initiated atom transfer radical polymerization in water or phosphate buffered saline. Incubation of uncoated NC with human serum albumin solution, fetal bovine serum, or human blood plasma resulted in fast aggregation observed by dynamic light scattering as an increase in diameter of particles present in the solutions. Conversely, these biological fluids affected only marginally the size distribution of the NC coated with a 60 nm thick poly(MeOEGMA) layer. The high suspension stability of the coated NC in complex biological fluids was related to the suppressed deposition of proteins from these fluids observed by surface plasmon resonance (SPR) on analogous poly(MeOEGMA) layer prepared on flat surfaces of SPR chips. © 2010 Elsevier B.V. All rights reserved.

1. Introduction A great deal of attention has been paid to the fabrication of nano-sized hollow polymeric particles due to their potential application in catalysis, separation, chromatography, diagnostics, drug delivery, and biomolecular-release systems [1–7]. Among them, self-assembled polymeric nanoparticles are ideal candidates for encapsulation of nonpolar guest molecules into their hydrophobic core [8–10]. A hydrolytically or enzymatically degradable polymer shell allows the guest molecules to be released in the surrounding biological medium. There is only a limited number of polymers that can be used as constituent of nanoparticles designed for in vivo applications [11–13]. The material of the NC should be degradable to non-toxic low-molecular mass products that can be easily excreted from the organism, e.g. polylactide and polycaprolactone [11–13]. Owing to their small sizes, nanoparticles are able to gain access to almost any tissue in the organism, and, if they are sufficiently small, to penetrate inside the cells [1,3,14]. In vivo biocompatibility of nanoparticles is mostly determined by

∗ Corresponding author at: Heyrovsky Sq. 2, 162 06, Prague, Czech Republic. Tel.: +420 296809234; fax: +420 296809410. E-mail addresses: [email protected], [email protected] (C. Rodriguez-Emmenegger).

the biophysicochemical characteristics of their surfaces [3,5,15]. Upon contact with biological media, their surface properties are rapidly changed by coating with proteins [16–21]. Subsequent colloidal instability [20,21] of the nanoparticles (e.g. particle aggregation, flocculation, precipitation, etc.) or adsorption of undesirable proteins can impair the designed particle functions and initiate unfavorable biological responses. Many of the early stage biological responses are determined by the nature of the deposited protein layer. Adsorbed proteins can affect the interaction of nanoparticles with cells and their behavior in the blood stream. When administrated intravenously, nanoparticles are rapidly cleared from the blood stream because they are recognized by cells of the mononuclear phagocytic system (MPS) or by the complement system due to opsonization, i.e. the binding of antibodies and/or complement molecules [22–25]. Nearly nothing is known about interaction of nanoparticles with blood coagulation factors and platelets by which hemostasis can be perturbed. Generally, nanoparticles with suppressed protein adsorption are of great interest to most applications in complex biological media. Currently, nanoparticles decorated with different antifouling polymers are subject of intense research [9,26–30]. Polymer coated ultra-low fouling gold, silica, and iron oxide nanoparticles have been prepared for imaging, diagnostics, and sensing [31–35]. However, their application as carriers of biologically active compounds is arguable [6,36,37]. Polymeric nanoparticles have been exten-

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Fig. 1. Synthesis of non-fouling nanocapsules. The surface initiated ATRP poly(MeOEGMA) shells resulted in particle diameters controlled within an interval of 170–900 nm.

sively studied as drug carriers [6,9,26,27,38]. Modifications aiming at increasing nanoparticle stability and prolonging their circulation in blood are of high importance for such systems [9]. Several authors have studied the interaction of polymeric nanoparticles with model protein solutions [25,39–41], while fouling properties of nanocarriers in complex biological media have not been sufficiently addressed yet. There has been a common tendency to misinterpret an observed decrease in adsorption from single protein solutions as an evidence of non-fouling properties in complex biological media [42–44]. However, the reduced or even totally suppressed non-specific adsorption from single protein solutions of serum albumin, IgG, fibrinogen, and lysozyme observed on various anti-fouling planar surfaces did not result in the suppressed fouling from blood serum and plasma [16]. Anti-fouling polymeric nanoparticles have been mostly prepared from different PEGcontaining polymers or by grafting PEG chains onto their surface [28,45,46]. Recent studies on “PEGylated” nanoparticles indicate that polymer architecture favoring higher surface concentrations of PEG results in a better anti-fouling ability [46]. Generally, a high density of attached polymers can be achieved by a grafting polymer chains from the surface. Atom transfer radical polymerization initiated from the nanoparticle surface is presented in this work as a novel organic-solvent-free technique for coating nanoparticles in aqueous solutions with non-fouling polymer brushes. Using this technique nanocapsules from FDA-approved biodegradable poly(␧caprolactone) were coated under mild conditions with non-fouling mono-methoxy terminated oligo(ethylene glycol) methacrylate polymer brush layers with the thickness controlled at a nanometer scale up to 350 nm (Fig. 1). The coating prevented the nanocapsules from aggregation in albumin solution, fetal bovine serum, and human blood plasma.

2. Experimental 2.1. Materials All chemical reagents were used without further purification. Mono-methoxy terminated oligo(ethylene glycol) methacrylate, Mn 300 (MeOEGMA), ␣,␻ (hydroxyl) poly(␧caprolactone), Mw = 65,000 Da (PCL), Span 60® , ␣-bromo isobutyric acid, 98%, N-hydroxysuccinimide, 98%(NHS), N-ethyl-N -(3diethylaminopropyl) carbodiimide hydrochloride, 99% (EDC), CuBr2 (99.999%), 2,2 -dipyridyl (99%), CuCl (99.995% trace metal

basis), human serum albumin, 99% (HSA), fetal bovine serum, non-USA origin (FBS), pooled blood plasma (HBP), and phosphate buffered saline (PBS) were purchased from Sigma–Aldrich. Acetone, Mygliol 812® oil, and cellulose membrane, 6,000–8,000 Da Spectra/Pore® were purchased from Lach-Ner, s.r.o., Sasol, and SpectrumLabs, respectively.

2.2. Preparation of PCL nanocapsules Polymeric nanocapsules (NC) of ␣,␻ hydroxy poly(␧caprolactone) (PCL) containing Mygliol 812® in the core were prepared by interfacial polymer deposition following solvent displacement according to the methodology described elsewhere [47,48]. PCL, Mygliol® and Span 60® were dissolved in acetone and subsequently injected over the aqueous phase. Acetone was evaporated and the suspension was concentrated under reduced pressure. Initiator decorated nanoparticles: A solution of ␣-bromo isobutyric acid, 0.15 M, N-hydroxysuccinimide, 0.05 M, and N-ethylN -(3-diethylaminopropyl) carbodiimide hydrochloride, 0.2 M, in MilliQ water was kept for 7 min at 25 ◦ C to activate the carboxylic groups and then transferred to the NC suspension, 5.5 × 109 mL−1 (determined by UV-turbidimetry) [47]. After stirring overnight, the solution was dialyzed for 12 h using a cellulose membrane Spectra/Pore® (6,000–8,000 Da). (All at room temperature.) Modification of the NC surface with poly(MeOEGMA): A solution of CuBr2 (8.1 mg, 36.4 ␮mol), 2,2 -dipyridyl (145 mg, 930 ␮mol), and MeOEGMA (5.7 g, 19 mmol) in 10 mL of water or PBS was degassed by Ar bubbling for 1 h. CuCl (37 mg, 375 ␮mol) was added under Ar atmosphere and the Ar bubbling continued for 30 min. The solution was added to 5 mL of the initiator decorated NC dispersion and the polymerization was allowed to proceed at 30 ◦ C under Ar. One milliliter samples were collected after different polymerization times. (All at room temperature) Physicochemical characterization of the NCs: Size distributions of the NC and NC coated with poly(MeOEGMA) (NC@MeOEGMA) were determined by quasi elastic light scattering (QELS) at an angle of 173◦ and at 25 ◦ C using a Zetasizer® ZS (Nanoseries, Malvern, UK). Mean size distribution values were obtained from 3 independent measurements. Measurements of the ␨-potential in MilliQ water were performed with the same instrument. Ten measurements were carried out to check the reproducibility. The measurements of the electrophoretic mobility were converted to ␨-potential (mV) using the Smoluchowski approximation. Transmission electron

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microscopy (TEM) was carried out using a microscope JEM200CX (Jeol, Japan). The NCs were deposited onto a copper TEM grid (300 mesh) coated with a carbon film. The images were taken at an acceleration voltage of 100 kV and recorded with a digital camera. Preparation of model planar surfaces: Silicone wafers or glass plates were coated by gold films deposited in vacuum for ellipsometry or surface plasmon resonance measurements, respectively. PCL was spin-coated on the gold surface from chloroform solution. Poly(MeOEGMA) brushes were grafted from a selfassembled monolayer of ␻-mercapto undecylbromoisobutyrate initiator attached to the gold surface using surface initiated atom transfer radical polymerization (ATRP). The polymerization was carried out at 30 ◦ C for 40 min in the same way as described above for the NC. Thickness of both the poly(MeOEGMA) layer and the spin-coated PCL film was 40 nm as determined by ellipsometry. Spectroscopic ellipsometry: The measurements were performed using Variable Angle Spectroscopic Imaging Auto-Nulling Ellipsometer EP3 -SE (Nanofilm Technologies GmbH, Germany) in the wavelength range of  = 398.9–811 nm (source Xe-arc lamp, wavelength step ∼10 nm). The measurements were performed at angle of incidence AOI = 70◦ in air at room temperature. For assessing the uniformity of the modifications, all samples were measured at three points. The obtained SE spectra were fitted with multilayer models in the EP3 -SE analysis software. The thickness and refractive indexes of polymer layers were obtained from simultaneous fitting of the obtained ellipsometric data using Cauchy relationship model. Due to low penetration depth of light, the Au layer was modeled as bulk gold using predefined EP3 -SE dispersion function. Surface plasmon resonance (SPR): A custom-built SPR instrument based on the Kretschmann geometry of the attenuated total reflection method and spectral interrogation of the SPR conditions from the Institute of Photonics and Electronics, Academy of Sciences of the Czech Republic was used. SPR chips were prepared by vacuum deposition of a gold layer (thickness approximately 50 nm) onto glass slide coated with an adhesion titanium layer (thickness approximately 2 nm). The tested solutions (HSA, 5 mg mL−1 , FBS, 10 and 100%) were driven by a peristaltic pump through four independent channels of a flow cell in which SPR responses were simultaneously measured. The non-specific protein deposition was observed as a shift in the resonant wavelength, res . Irreversible protein deposition after 15 min of incubation with the biological fluid was determined from the difference in resonant wavelength, res , measured in PBS before and after the incubation. A shift re = 1 nm detected by SPR was related to a deposited protein mass of 200 pg mm−2 . The relationship was estimated by model experiments in which HSA adsorption was observed using SPR and the deposited HSA mass was determined by FTIR GASR (see below). The same correlation was used for other proteins assuming a similar refractive index. The limit of the SPR detection defined as a three times the sensor response of the standard deviation of the baseline noise was determined to be res = 0.03 nm, which roughly corresponded to 6 pg of deposited proteins per mm2 . The protein mass in a deposit was estimated by comparing res with that of the adsorbed HSA monolayer used for FTIR GASR calibration described below. FTIR grazing angle specular reflectance (FTIR GASR) calibration: A calibration curve was obtained by spreading various amounts of HSA dissolved in water over gold surface of SPR chips, drying them and measuring their FTIR GASR spectra with a FTIR spectrometer Bruker IFS 55 equipped with Pike Technologies 80Spec GASR attachment. HSA deposition was measured as the re in SPR chips. The mass of HSA adsorbed on these SPR chips after the incubation with HSA solution, washed with PBS, and dried was determined by comparing FTIR GASR spectra with the calibration curve mentioned above.

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Fig. 2. Kinetics of the surface-initiated ATRP of MeOEGMA brush layers on poly(PCL) nanocapsules in water (triangles) and PBS (circles). The layer thickness was determined by QELS.

3. Results and discussion Diameters of about 170 nm determined from the particle size distribution measured by QELS and ␨-potential of −23.3 ± 0.5 mV were characteristics of ␣,␻ (hydroxy) poly(␧-caprolactone) nanocapsules prepared by interfacial polymer deposition. The value of ␨-potential corresponds with those reported earlier [27]. The nanoparticles were stable in aqueous solutions probably due to the negative ␨-potential caused by charged carboxy-terminus of PCL chains exposed at NC surface. Hydroxyl groups of PCL present at the NC surface [49] made it possible to attach ␣-bromo isobutyric acid initiator activated by EDC/NHS. A shift in the ␨-potential to the less negative value of −13.6 ± 0.9 mV suggested that also some carboxyl groups on the NC surface took part in the initiator attachment. The MeOEGMA was grafted from the NC in water or PBS using ATRP initiated from ␣-bromoisobutyrate anchored to the NC surface. The growth of polymer brushes resulted in the increase in NC diameter observed by QELS. The size distributions of particles present in the samples collected at different polymerization times were determined by QELS. The thickness of the polymer layer growing on the NC surface was estimated by subtracting the mean radius of 85 nm of the initial NC calculated for the initial size distribution from the mean radius of the polymer coated NC calculated for the size distribution measured at the selected polymerization time. A linear increase in the polymer layer thickness with time (Fig. 2) indicated that the surface initiated ATRP of MeOEGMA was well-controlled. A lower rate of polymerization was observed when the solvent was PBS as expected. ATRP uses a catalytic amount of a transition metal complex, which reversibly abstracts a halogen atom from a polymer chain end, thereby transforming the latter into an active propagating radical from a dormant state. The high concentration of the extraneously added NaCl present in the buffer resulted in halogen exchange favoring the dormant chains capped with chlorine over those with bromine [50,51]. Due to the higher bond energy C–Cl than C–Br the equilibrium was shifted toward the dormant chains and consequently resulted in a drop in the rate of polymerization [50]. The observed shift in ␨-potential to less negative values with increasing polymerization time could be explained by the difference between the static dielectric permittivity of the hydrated brushes and of the ambient aqueous solution, as stated by the Coehn and Raydt phenomenological rule [52] Coated NC characterized by ␨-potential of −6.5 ± 0.5 mV and a mean diameter of 290 nm, i.e. NC coated with a polymer layer thick 60 nm, were chosen for testing

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Fig. 4. Interaction of NC@MeOEGMA (a) and uncoated NC (b) with solution of human serum albumin (HSA), fetal bovine serum (FBS), and undiluted human blood plasma (HBP). Diameter distribution of NC@MeOEGMA (a) and uncoated NCs (b) was measured by QELS in PBS (black) and after incubation with HSA solution, 5 mg mL−1 PBS (grey dotted), FBS 10% PBS (black dotted), or HBP (grey) for 15 min. (The black-dotted curve for FBS superimposes the grey one for HBP in (b).)

Fig. 3. TEM images of NC coated with poly(MeOEGMA) brush prepared by ATRP in PBS for 90 min and dried (A and B). A dark initial NC of a diameter about 170 nm is surrounded by a grey corona of the poly(MeOEGMA) coating (B).

in biological fluids. While the suspension of uncoated NC with ␨potential of −23.3 ± 0.5 mV was probably stabilized by the electric repulsion, the suspension of coated NC with much less negative ␨-potential was obviously stabilized due to a steric effect of the poly(MeOEGMA) shell [53]. Transmission electron microscopy of the dried samples showed that all particles present in the suspension were of spherical shapes. A TEM micrograph of typical nanoparticles is shown in Fig. 3. A dark initial NC of a diameter about 170 nm is surrounded by a grey corona of the poly(MeOEGMA) coating (Fig. 3B). Changes in the size distribution of uncoated NC and NC coated with a 60 nm MeOEGMA shell induced by the addition of HSA solution, 5 mg mL−1 PBS, fetal bovine serum, 10% in PBS, or undiluted human blood plasma were observed by QELS (Fig. 4). It should be noted that all the QELS measurement were carried out without any sample pretreatment, such as, the frequently used centrifugation, to avoid the removal of any big aggregates. The size distributions of particles reached steady state values shown in Fig. 4 in time periods shorter than 1 min after adding the biological fluids. It was observed that upon addition of HSA, FBS or HBP the size distribution of NC@MeOEGMA with a mean diameter value of 290 nm changed only marginally (Fig. 4a) while the size distribution of uncoated NC with a mean diameter of 170 nm was shifted to markedly larger diameters reaching mean values of about 470 nm in HSA and 320 nm in FBS or HBP (Fig. 4b). The capability of the poly(MeOEGMA) brushes to prevent protein deposition from biological fluids was tested on model planar surfaces prepared on SPR chips. The amounts of proteins

deposited at the PCL surface after 15 min incubation with solutions of HSA, 5 mg mL−1 PBS, 10% FBS in PBS, and 100% FBS were 760 ± 90 pg mm−2 , 1000 ± 23 pg mm−2 , and 1050 ± 92 pg mm−2 in HSA, 10% FBS, and 100% FBS, respectively. The results indicated that the deposits did not exceed a monomolecular layer of proteins. For example, a monolayer of closely packed HSA molecules with native conformation can be theoretically estimated from the HSA molecular dimensions to contain 9000 pg mm−2 HSA in a 6 nm thick layer. There was no deposition from these biological fluids on the poly(MeOEGMA) brushes (Fig. 5) observable within the limit of our SPR detection of about 6 pg of a protein deposit per mm2 . The observed big increase in size of particles in the suspension of uncoated NC from a mean diameter of 170 nm in PBS to 470 nm in HSA and 320 nm in FBS was probably not due the formation of a thick protein corona around individual particles but rather due to the NC clustering induced by protein adsorption at the surface of uncoated PCL shell. A dramatic increase in the effective “size” of gold nanoparticles due to their aggregation mediated by adsorption from single protein solutions of HSA, and other common blood plasma proteins was described and discussed by Lacerda et al. [54]. Marginal changes in the size distribution of NC@MeOEGMA after

Fig. 5. Protein deposits irreversibly adsorbed from HSA solution and fetal bovine serum on surfaces of spin coated PCL film and MeOEGMA brushes prepared on SPR chips. The surfaces were exposed to flowing solutions of HSA, 5 mg mL−1 in PBS and fetal bovine serum 10% and 100% for 15 min. The standard deviations were calculated from three independent measurements. Protein adsorption on poly(MeOEGMA) was below the detection limit of our SPR (6 pg mm−2 ).

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adding protein solutions suggest that similarly to poly(MeOEGMA) brushes on flat surfaces, poly(MeOEGMA) shell prevents also NC from fouling. Such property indicates the potential of this modification to protect nanocarriers from undesirable interactions in biological fluids and, perhaps, even in physiological environments. 4. Conclusions Poly(MeOEGMA) brushes of up to 350 nm were grafted from the surface of nanocapsules prepared from of FDA-approved biodegradable poly(␧-caprolactone). This is the first time that wellcontrolled polymerization of a shell was successfully achieved from the surface of polymeric nanocapsules in water and PBS. The procedure was completely carried out in aqueous solutions. Thus, it could be applied to various colloids that are not stable in the presence of organic solvents. The negligible protein adsorption with the brushes prepared on model flat surfaces and the resistance of the coated nanocapsule suspension to the effect of human serum albumin solution, fetal bovine serum, and undiluted human blood plasma indicate that this surface modification can be applied for the preparation of potential carriers of bioactive compounds in complex biological media. Acknowledgments This research was supported by the Academy of Sciences of the Czech Republic under contract No.KAN200670701 and by the Grant Agency of the Czech Republic under contract No.P208/10/1600. ˇ Dr. Miroslav Slouf from Institute of Macromolecular Chemistry AS CR is acknowledge for the TEM images. A. Jäger, E. Jäger, A. R. Pohlmann and S.S. Guterres would like to thanks CNPq/MCT/Brazil and FAPERGS/RS/Brazil. References [1] H. Hillaireau, P. Couvreur, Nanocarriers’ entry into the cell: relevance to drug delivery, Cell. Mol. Life Sci. 66 (2009) 2873–2896. [2] C. Boyer, M.R. Whittaker, C. Nouvel, T.P. Davis, Synthesis of hollow polymer nanocapsules exploiting gold nanoparticles as sacrificial templates, Macromolecules (2010). [3] A.E. Nel, L. Madler, D. Velegol, T. Xia, E.M.V. Hoek, P. Somasundaran, F. Klaessig, V. Castranova, M. Thompson, Understanding biophysicochemical interactions at the nano-bio interface, Nat. Mater. 8 (2009) 543–557. [4] M.C. Daniel, D. Astruc, Gold nanoparticles: assembly, supramolecular chemistry, quantum-size-related properties, and applications toward biology, catalysis, and nanotechnology, Chem. Rev. 104 (2004) 293–346. [5] I. Lynch, K.A. Dawson, Protein–nanoparticle interactions, Nano Today 3 (2008) 40–47. [6] C. Vauthier, K. Bouchemal, Methods for the preparation and manufacture of polymeric nanoparticles, Pharm. Res. 26 (2009) 1025–1058. [7] S. Lee, J.H. Ryu, K. Park, A. Lee, S.-Y. Lee, I.-C. Youn, C.-H. Ahn, S.M. Yoon, S.-J. Myung, D.H. Moon, X. Chen, K. Choi, I.C. Kwon, K. Kim, Polymeric nanoparticlebased activatable near-infrared nanosensor for protease determination in vivo, Nano Lett. 9 (2009) 4412–4416. [8] Y. Bae, K. Kataoka, Intelligent polymeric micelles from functional poly(ethylene glycol)–poly(amino acid) block copolymers, Adv. Drug Deliv. Rev. 61 (2009) 768–784. [9] V.P. Torchilin, Multifunctional nanocarriers, Adv. Drug Deliv. Rev. 58 (2006) 1532–1555. [10] A. Boker, J. He, T. Emrick, T.P. Russell, Self-assembly of nanoparticles at interfaces, Soft Matter 3 (2007) 1231–1248. [11] J.S. Chawla, M.M. Amiji, Biodegradable poly(epsilon-caprolactone) nanoparticles for tumor-targeted delivery of tamoxifen, Int. J. Pharm. 249 (2002) 127–138. [12] D. Lemoine, C. Francois, F. Kedzierewicz, W. Preat, M. Hoffman, P. Maincent, Stability study of nanoparticles of poly(epsilon-caprolactone), poly(d,l-lactide) and poly(d,l-lactide-co-glycolide), Biomaterials 17 (1996) 2191–2197. [13] R. Gref, P. Couvreur, G. Barratt, E. Mysiakine, Surface-engineered nanoparticles for multiple ligand coupling, Biomaterials 24 (2003) 4529–4537. [14] T. Cedervall, I. Lynch, S. Lindman, T. Berggard, E. Thulin, H. Nilsson, K.A. Dawson, S. Linse, Understanding the nanoparticle-protein corona using methods to quantify exchange rates and affinities of proteins for nanoparticles, Proc. Natl. Acad. Sci. U. S. A. 104 (2007) 2050–2055. [15] I. Lynch, A. Salvati, K.A. Dawson, Protein–nanoparticle interactions. What does the cell see? Nat. Nanotechnol. 4 (2009) 546–547.

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