Elastic Properties of a Protein–Polymer-Grafted Surface - American

Article pubs.acs.org/Langmuir

Elastic Properties of a Protein−Polymer-Grafted Surface Kristen Keller,† Armen Amirian,‡ and Pinar Akcora*,† †

Department of Chemical Engineering and Materials Science, and ‡Department of Mechanical Engineering, Stevens Institute of Technology, Castle Point on Hudson, Hoboken, New Jersey 07030, United States S Supporting Information *

ABSTRACT: Surfaces grafted with poly(methyl methacrylate) (PMMA) and streptavidin were synthesized through click chemistry to investigate the role of surface stiffness on protein adsorption as the hydrophilic and hydrophobic surface coverage of the substituents vary. Surface topographies coupled with the nanoindentation results indicated that, with the appropriate selections of polymer coverage and chain length, the extent of non-specific protein adhesion could be controlled by the hydrophobic interactions between PMMA, biotin, and streptavidin. It was shown that, when the molecular weight and stiffness of PMMA was close to that of streptavidin, patchy PMMA morphologies were obtained, which help inhibit the nonspecific adsorption of streptavidin.

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INTRODUCTION It is of great importance to understand the effects of surface chemistry, topography, and elasticity on protein adhesion when designing biomedical in vivo devices. Non-specific protein adhesion, which occurs when a synthetic material is exposed to a biological fluid, can trigger a variety of undesirable effects, such as inflammation and implant failure. Roughness, hydrophobicity, and elastic modulus of the exposed surface all influence the extent to which repulsion or attraction between two contacting materials will occur. Therefore, it is critical to understand how these properties will affect protein adhesion when designing any biocompatible surface, such as implants, drug carriers, and biosensors. Polymer brushes are commonly used to help control such adhesion between the surface of a synthetic material and biological entities in the surrounding medium.1 Polymer grafting density, chain length, polarity, and surface hydrophobocity/hydrophilicity all play an important role in nonfouling and protein immobilization on polymerdecorated surfaces.1−9 To date, films consisting of highly hydrophilic polymers, such as poly(ethylene oxide) (PEO), poly(ethylene glycol) (PEG), and oligo(ethylene glycol) precursors, have had the most success in preventing protein adhesion.3,10−12 Here, we demonstrate a novel heterogeneous surface consisting of both protein and biocompatible polymer brushes, which exhibit patchy morphologies, and further study their elastic properties before and after attaching proteins. Analysis of the mechanical properties of our surfaces ultimately will help us to evaluate how surface elasticity, a factor that affects adhesiveness, changes when proteins are adsorbed onto a hydrophobic surface. Topological and mechanical characterization of these functionalized surfaces can lead to new assembly routes of nanoparticles for the design of novel complex nanomaterials. © 2012 American Chemical Society

Surfaces of controlled roughness and hydrophobicity have been created using techniques such as lithography, nanopatterning, preparation of mixed brushes, and nanoparticles embedded in temperature-sensitive polymers. Lithography provides an efficient means to create nanoscale patterns on a surface layer, thus manipulating the overall surface roughness and protein adsorption.13 Nanopatterning of biomolecules on functionalized surfaces is similarly used to immobilize proteins and to prepare the protein nanoarrays.14 Additionally, nanoparticles within temperature-responsive polymers are shown to control adhesive properties of a surface to various materials in aqueous medium, where particle size and length of responsive polymer chains determine the low adhesiveness.15 Binary mixed brush systems have also been designed for biomedical use in attempts to minimize protein adhesion.7,16−19 Here, polymer constituents respond to changes in the external environment, making it possible to effectively switch protein adhesion on and off, for example, by switching from a hydrophobic surface to a hydrophilic surface.7,20−22 To our knowledge, surfaces of varying hydrophilicity upon integrating two classes of macromolecules, both synthetic and biological, to form mixed brushes have not been designed to inhibit protein adsorption. One popular and universal method of fabricating various brush surfaces is click chemistry, which has been used to graft both polymer chains and biological molecules onto planar substrates as well as the curved surfaces of nanoparticles, carbon nanoparticles, and nanofibers.23−28 For example, the versatility and efficiency of click chemistry for grafting different polymers with the same reaction was demonstrated when azidefunctionalized PEG, poly(methyl methacrylate) (PMMA), and Received: September 22, 2011 Revised: January 12, 2012 Published: January 24, 2012 3807

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Attachment of Acetylene-Functionalized Biotin. Azidefunctionalized silicon wafers were each immersed in a solution consisting of acetylene-functionalized biotin (synthesized according to a previously reported procedure) in phosphate-buffered saline (PBS) (0.606 mL), CuSO4 (3 mg, 0.019 mmol), TEA (0.211 g, 2.09 mmol), sodium ascorbate (21 mg, 0.11 mmol), methanol (4.06 mL), and toluene (1.03 mL).29 The solutions were sonicated for 1 h at room temperature and washed thoroughly in toluene. PMMA Grafting. Biotin-functionalized silicon wafers were each immersed in a 50 mg/mL solution of acetylene-functionalized PMMA, which was synthesized following the reported protocol.32 CuBr (0.003 g, 0.021 mmol) and N,N,N′,N′,N″-pentamethyldiethylenetriamine (PMDETA) (4.4 μL, 0.021 mmol) were added to the solution, and the reaction flask was purged with N2 for 20 min.33 The sealed flask was placed in an oil bath at 60 °C and allowed to react for 24 h. Samples were thoroughly washed in toluene and annealed for 8 h under reduced pressure at 100 °C. Streptavidin Adsorption. One side of the Si wafer was treated with a 10 μg/mL solution of streptavidin in PBS (100 μL) for 1 h. The wafer was rinsed thoroughly in deionized (DI) H2O and dried under a stream of N2. Wafers were then rinsed 4 times in a 0.05% solution of Tween 20 in PBS under constant, vigorous shaking for 1 min per rinse. This was followed by four rinses in PBS under similar agitation conditions.34 Finally, samples were rinsed with DI water, blow-dried under filtered nitrogen, and annealed at room temperature for 24 h under reduced pressure. Fourier Transfer Infrared (FTIR) Measurements. Transmission infrared (IR) spectra were collected on a Bruker Instruments’ Tensor 27 FTIR spectrometer using a resolution of 8 cm−1 and a total of 1000 scans per sample. For all transmittance measurements of the functionalized silicon surfaces, the initially cleaned substrates were used as a reference. Atomic Force Microscopy (AFM) Measurements. AFM measurements were performed in air using a NanoInk NSCRIPTOR Dip Pen Nanolithography system. Silicon AFM probes with spring constants (kc) of 40 N/m and resonance frequencies of 300 kHz were purchased from Vista Probes. InkCAD software was used to obtain AFM images. Images were obtained in tapping mode with a scan rate of 0.8 Hz. Force−distance measurements for the Young’s modulus calculation were obtained in contact mode using SPM Cockpit software. Each measurement was taken over a 1.02 s interval. Force− distance measurements were performed on a straight line, typically consisting of 10−25 measurements taken at step sizes of 3 μm apart. A total of 100 measurements were recorded for every sample. Indentation profiles are shown in the Supporting Information. The Young’s modulus (Ei) is calculated according to the Hertz model:35 Et = 3/4(1 − ν2)kcd/R1/2h3/2, where d is the deflection, h is the depth of indentation, and kc is the cantilever spring constant. Poisson’s ratio, ν, of 0.4 and the radius of the tip, R, of ∼10 nm (on the basis of the specifications of the manufacturer) are used. Contact Angle Measurements. The contact angle of a water droplet on surfaces was measured using a KSV Instruments LTD CAM101 system.

polystyrene (PS) were clicked to alkyne-functionalized silane monolayers on silicon wafers.23 In the present work, we have developed a heterogeneous, biocompatible surface by grafting the hydrophilic protein, streptavidin, and hydrophobic biocompatible polymer, PMMA, with molecular weights of 49 360 g mol−1 [polydispersity index (PDI) = 1.08] and 14 950 g mol−1 (PDI = 1.13), onto silicon substrates through click chemistry. These unique surfaces were fabricated in a stepwise fashion, shown schematically in Figure 1A. The acetylene-

Figure 1. (A) Schematics of stepwise fabrication of streptavidin− PMMA-tethered heterogeneous silicon surfaces. (B) Transmittance IR spectra of surfaces functionalized with (a) APS, (b) azide (N3), (c) biotin, and (d) biotin−PMMA. (Inset) Evolution of the azide band at ∼2100 cm−1.

functionalized biotin (which promoted streptavidin adsorption in a later step) and PMMA were immobilized onto azidefunctionalized silicon substrates in two separate reactions.24,28,29 PMMA- and biotin-functionalized surfaces were then immersed in streptavidin solution and rinsed rigorously with water. Protein attachment was simply facilitated by the strong affinity between biotin and streptavidin.

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EXPERIMENTAL SECTION

Materials. Toluene [high-performance liquid chromatography (HPLC) grade], methanol, and ethanol were purchased from PHARMCO-AAPER. All other chemicals were purchased from Sigma-Aldrich. Methyl methacrylate (MMA), toluene (for PMMA synthesis), and CH2Cl2 were distilled over CaH2 before use. All other chemicals were used as received. N-Doped (100)-oriented silicon wafers with thicknesses of 356−406 μm and 1−10 Ω cm resistivities were purchased from Silicon Quest International. Aminopropylsilane (APS) Monolayer Formation. The hydrophilic silicon wafers were each placed into separate vials containing 10 mL of toluene preheated to ∼70 °C. Under a N2 atmosphere, 3(aminopropyl)triethoxysilane (APTES) (0.0427 mmol) was added to each vial and the wafers were incubated for a period of 7.5 h.30 Wafers were then washed with toluene, dried under filtered N2, and sonicated in toluene for 15 min. Surface Diazo Transfer to Amines. Triflyl azide was prepared according to a previously reported procedure.31 Si−APS substrates were sonicated for 30 min in solutions of triflyl azide (1.03 mL), CuSO4 (3 mg, 0.019 mmol), triethylamine (TEA) (0.211 g, 2.09 mmol), H2O (0.606 mL), and methanol (4.06 mL). Azidefunctionalized silicon substrates (Si−N3) were then thoroughly washed in toluene and dried under a stream of filtered nitrogen.

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RESULTS AND DISCUSSION The changes in the chemical composition of the silicon surfaces after each step of the surface functionalization were monitored through FTIR measurements (Figure 1B). Specifically, the evolution of the characteristic azide peak at ∼2100 cm−1 was the most important indicator of the success of the surface modifications via click reactions. The appearance of the characteristic azide peak at ∼2100 cm−1 is first observed when an aminylated silicon substrate (fabricated through deposition of an APS monolayer) is modified with surface azide moieties. Subsequently, when biotin and PMMA click to the azide sites, the amount of free azide groups at the substrate surface is reduced and a gradual decrease in the azide peak in the IR spectra is observed (shown in the inset of Figure 1B). It 3808

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was found to be significantly lower (3293), indicating that less azide sites are available for PMMA grafting. Surfaces with varying amounts of azide sites were further functionalized sequentially by clicking PMMA and immersing them into streptavidin solution. Figure 3 summarizes the surfaces that were synthesized and discussed in this work. Hydrophobic polymers clicked onto hydrophilic surfaces form patchy morphologies with their observed three-dimensional topographies shown in Figure 4A, and as streptavidin was adsorbed to the biotin sites, non-specific adsorption of streptavidin onto the PMMA regions of these heterogeneous surfaces was not intended. The morphology of the PMMA−streptavidin-functionalized surfaces was characterized in AFM. It is clear from Figure 4A that the pure PMMA brush forms homogeneous globular-like morphologies, while a highly heterogeneous morphology is observed for all of the surfaces where both biotin and PMMA are attached to the azide sites. Here, sharp interfaces exist between patchy areas of tethered PMMA and areas of lower topographies, presumed to be regions of attached biotin. The repulsion of the long hydrophobic PMMA chains from hydrophilic biotin is presumably negligible. Although the PMMA coverage within the three biotin−PMMA-tethered surfaces is changing, similar features are observed in the images of each surface. When these surfaces are incubated in streptavidin solutions (Figure 4B), globular structures within the PMMA patches appear to be more compact than those observed in Figure 4A, while regions of lower topography are still apparent. We conjecture that PMMA and streptavidin will repel each other because of the differences in hydrophilicity of the two components. An entropic constraint will therefore be imposed on the PMMA chains in the vicinity of streptavidin, causing the PMMA patches to be more compact. The fluctuation of height profiles is noted after adding streptavidin (see Figure S1 of the Supporting Information). We believe that anchoring a hydrophilic component (streptavidin) and washing surfaces in water will enhance the phase separation of polymer and result in surfaces where PMMA regions are higher. The pure PMMA brush washed with streptavidin also displayed similar globule features; however, it is inconclusive from surface topography images if the protein has adhered onto pure PMMA (Figure 4B). The enlarged features of the 0.01 mg/mL sample before and after anchoring streptavidin in Figure S2 of the Supporting Information clearly show the phase-separated polymer grafts after streptavidin attachment. Previous work has shown that streptavidin can be immobilized on patterned surfaces of PMMA, especially at the edges where the surface energy is high;41−43 therefore, it is expected that streptavidin will be non-specifically adsorbed to PMMA, an effect that will be rationalized through nanoindentation measurements. To reveal how streptavidin and PMMA would interact when immobilized to the same surface, the Young modulus along the heterogeneous surfaces was calculated through the Hertz model using data from nanoindentation profiles, which were obtained through force−distance measurements in AFM. Samples were analyzed by taking force−distance measurements along a line across the samples at step sizes of 3 μm, with ∼100 measurements per sample. When these results were compared to the rigid silicon substrate, we could deduce the measurements that represented the softer areas of PMMA or streptavidin along a sample (see Figures S6−S8 of the Supporting Information for nanoindendation profiles). The Young modulus for each of these measurements was then

is well-known that the grafting density of polymer chains is an important parameter to control protein adsorption;36−38 therefore, we designed polymer-tethered surfaces with varying coverages. To vary the surface coverages of biotin and PMMA, azide-functionalized surfaces were treated with biotin concentrations of 0.1, 0.01, and 0.001 mg/mL during the initial click reaction, which in turn limited the number of remaining surface azide functionalities available for PMMA grafting in the second click reaction. The reaction efficiency upon changing the biotin concentration was also assessed through FTIR measurements by integrating the area under the azide peak before and after biotin attachment. A 90% disappearance of the azide peak was observed with biotin concentrations of 0.1 mg/mL, while a 55% disappearance was found for biotin concentrations of 0.001 mg/mL. It is important to note that our focus is not to determine the exact grafting densities because we do not measure the amounts of adsorbed proteins, but rather we focus on the change of stiffness of a biocompatible surface when streptavidin is physically associated to surfaces, through either non-specific or specific adsorption to biotin. The local grafting density of pure PMMA patches was calculated to be ∼0.2 chains/nm2 according to de Gennes’ flat brush theory,39 using the thickness values measured in ellipsometry. X-ray photoelectron spectroscopy (XPS) experiments were conducted in attempts to quantify the coverage of biotin after it was clicked onto the Si−N3 surfaces. XPS spectra for a Si−N3 sample as well as Si−biotin samples of different biotin coverages are shown in Figure 2. For the Si−N3 sample, two

Figure 2. X-ray photoelectron spectra of N3- and biotin-functionalized surfaces.

peaks at 400 and 405 eV, corresponding to two oxidation states of the azide nitrogens, are observed.40 After the click reaction between Si−N3 and acetylene-functionalized biotin, a peak at 400 eV should remain for the triazole. It is observed in Figure 2 that the XPS peak shifts from ∼399 eV for samples having high and low biotin coverages to ∼401 eV for the Si−N3 and Si− biotin sample of an intermediate coverage. Analyses of these samples performed at different times can be the cause for this minor shift. We then integrated the areas under the XPS peak at ∼400 eV. The area of the Si−N3 sample was much larger (5676) than the areas of samples after biotin had been attached. Integrated areas for the samples of intermediate and low biotin coverages did not differ greatly (∼4800), indicating that these samples may have similar amounts of remaining azide. On the other hand, the area of the sample with a high biotin coverage 3809

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Figure 3. Schematic representation of hydrophilic silicon surfaces functionalized with long (49 kg mol−1) and short (15 kg mol−1) hydrophobic polymer brushes. Non-specific protein adsorption is prevented by the patchy topography of the surface, which is formed by the 49 kg mol−1 brush.

Figure 4. AFM topography images of 49 kg mol−1 PMMA grafted to Si substrates (A) before and (B) after streptavidin attachment. The concentration of biotin used during the initial click reaction is indicated on the top. The first image in the bottom panel is the pure PMMA brush after incubation in a streptavidin solution.

points and shapes of the force curves of functionalized surfaces were compared to the data obtained for an unmodified silicon substrate, we were able to deduce the soft regions and rigid areas of the substrate. Measurements taken on rigid areas of biotin were already eliminated from these graphs. We then calculated the Young modulus using the Hertz model for the softer portions of the samples. The moduli measured at different positions along the sample are averaged. Thereby, the large deviations in moduli reflect the high heterogeneity of surfaces (Figure 5).

compared at the point where the force exerted between the AFM tip and sample was 2 μN, a force found within the linear region of the force−indentation data (see Figure S7 of the Supporting Information). Analysis of force−distance data obtained through AFM is quite complicated because force−distance measurements are recorded along a heterogeneous surface. Therefore, measurements will represent the force between the tip and either the biotin layer, PMMA grafts, or adsorbed streptavidin. Figure S5 of the Supporting Information shows the retract-approach force curves with respect to piezo displacement. When the deflection 3810

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Figure 5. Variation in Young’s modulus of (A) streptavidin washed pure PMMA brush (49.3 kg mol−1), silicon surface functionalized with streptavidin, pure PMMA brush, and PMMA−streptavidin immobilized surface, where a biotin concentration of 0.001 mg/mL is reacted with the surface. (B) Biotin and PMMA (49.3 kg mol−1) immobilized surfaces before and after streptavidin attachment. (C) AFM topography images of PMMA (14.9 kg mol−1)−biotin (0.001 mg/mL) surface before (left image) and after (right image) streptavidin attachment. (D) Young’s moduli of biotin and PMMA (14.9 kg mol−1) immobilized surfaces before and after streptavidin attachment. Values in panels A, B, and D were at the point of the measurement where the force between the AFM tip and the sample was 2 μN.

that the surface with the highest biotin coverage (i.e., the lowest PMMA coverage) is the stiffest compared to other biotin concentrations. We surmise that PMMA chains shrink to avoid unfavorable interactions with the water-retaining biotin surface, which becomes apparent at the highest biotin coverage. Consequently, PMMA limits its contact area with the biotinylated surface and becomes highly resistant to compression. Figure 5B also compares the elastic moduli of each biotinand PMMA-grafted surfaces before and after streptavidin attachment. It was observed that the addition of the protein does not have much effect on the stiffness of the surface. While this result indicates that the non-specific adsorption is minimal on the heterogeneous surfaces, it also supports the result concluded from AFM images that the distinct features of the observed patches are from segregated PMMA chains and not from non-specifically adsorbed proteins. Surfaces functionalized with shorter PMMA chain length (14.9 kg mol−1) showed more homogeneous features unlike the patchy morphologies (Figure 5C) of the 49.3 kg mol−1 PMMA. Additionally, protein attachment did not appear to change the topography (for more AFM images of surfaces with different polymer coverages, see Figures S3 and S4 of the Supporting Information). After streptavidin adsorption, surface stiffness increased slightly, which is attributed to the difference in

We first focus on the pure PMMA brush, which shows that uniform globular graft regions have an elastic modulus lower than 1 GPa with small variation of the data (Figure 5A). A 0.1 mg/mL biotin surface had features of a 12 nm height (see Figure S1 of the Supporting Information) and had the largest deviation in Young’s modulus data, as represented in Figure 5B. The surfaces reacted with 0.01 and 0.001 mg/mL biotin display ∼10 nm height features; however, their averaged moduli were below 1 GPa. The 0.001 mg/mL sample has a modulus value that does not deviate in different areas and is also comparable to the modulus of the pure PMMA, which indicates that the 0.001 mg/mL sample has more coverage of PMMA grafts than the other surfaces. We measured that the protein-washed pure PMMA brush has a higher Young’s modulus than both the pure PMMA brush and pure streptavidin attached to a biotinylated surface, indicating that streptavidin is physically adsorbed to the homogeneous PMMA brush (Figure 5A). We conjecture that the increase in the modulus is arising from the contribution of the protein behaving more stiffly on a hydrophobic polymer surface. Young’s moduli calculations of the brushes having different biotin coverages show that the elastic modulus decreases with an increasing PMMA coverage. Results obtained for each surface of varying biotin coverages (Figure 5B) reveal 3811

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Figure 6. Contact angle measurements on functionalized surfaces before and after streptavidin attachment for two different brush cases: (A) 49.3 and (B) 14.9 kg mol−1.

molecular weight of the protein (∼60 kg mol−1) and PMMA (14.9 kg mol−1) (Figure 5D), and their elastic modulus was found to be close to that of pure streptavidin (∼1.5 GPa). This result suggests that streptavidin can cover the short brushes (illustrated in Figure 3) and, hence, dominate the overall stiffness of the surface. We also tested the pure PMMA brush of 14.9 kg mol−1 after washing with streptavidin and found that its elastic modulus varies between 1 and 2 GPa, which is close to that of surfaces tethered with 14.9 kg mol−1 PMMA and streptavidin. This result is consistent with the result deduced from patchy morphologies of the 49 kg mol−1 brush that physical adsorption of protein becomes significant when the adsorbed surface is smooth (homogeneous), as in the short brush case. Additionally, we have measured the contact angle of a water droplet on our surfaces to determine the degree of hydrophobicity as PMMA coverage of the surfaces change (Figure 6). Pure PMMA is a hydrophobic surface, and when it is washed with streptavidin, its contact angle decreases to ∼80°, supporting our result that streptavidin is physically adsorbed on the pure PMMA-tethered surface. This result is consistent with the modulus data of non-specifically bound PMMA. The sample with the lowest PMMA coverage (treated with 0.1 mg/ mL biotin) may shrink to avoid interactions with biotin [as discussed in the Atomic Force Microscopy (AFM) Measurements section], therefore forming a slightly more hydrophobic surface than the two other functionalized surfaces. We observed that hydrophobicity of surfaces does not change significantly as the PMMA coverage increases, and it is not affected when the streptavidin is tethered for the 0.01 and 0.001 mg/mL samples. The latter result is interestingly in line with the modulus measurements, where moduli do not change with the streptavidin addition. Contact angle measurements support our discussion that, although the extent of hydrophocity does not change with the PMMA coverage, the surfaces also do not become more hydrophilic with streptavidin attachments, indicating that there is minimal non-specific protein adsorption on the patchy surfaces and streptavidin mainly binds to available biotin on the surface. We have also measured the contact angle of a water drop on the short brush (14.9 kg mol−1) system. As shown in Figure 6B, surfaces become more hydrophobic after adding streptavidin. At first sight, this might seem contradictory with the nanoindentation data, where the

stiffness did not change with the streptavidin coverage. However, contact angle results interestingly suggest that dewetting between PMMA chains and streptavidin may occur.

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CONCLUSION In conclusion, we have developed novel biocompatible surfaces functionalized by PMMA and a model protein (streptavidin). These surfaces exhibit highly heterogeneous, “patchy” morphologies, a factor that appears to limit the extent of nonspecific protein adhesion to the surface with the appropriate selections in polymer coverage and chain length. We have shown that nanoindentation measurements can be used to provide insight into protein adsorption on protein−polymer decorated surfaces. The elasticity of these heterogeneous surfaces was controlled with the hydrophobic interactions between the polymer, biotin functionalities, and protein. Nonspecific adsorption of streptavidin was found to be minimal on the patchy PMMA surfaces. Such biocompatible, patchy, polymer-grafted surfaces can find potential applications in biomedical devices, where inhibiting non-specific protein adhesion is critical. In addition, these heterogeneous surfaces can translate into nanoparticles functionalized with protein and polymer patches, which can lead to new self-assembly routes.

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ASSOCIATED CONTENT

S Supporting Information *

Topography images (Figures S1−S4), typical force versus piezo displacement plot (Figure S5), and force versus indentation profiles (Figures S6−S8). This material is available free of charge via the Internet at http://pubs.acs.org.

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AUTHOR INFORMATION

Corresponding Author

*Telephone: (201) 216-5060. Fax: (201) 216-8306. E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS We thank the Stevens Institute of Technology for start-up funds. We acknowledge the Multiscale Imaging Lab (LMSI) of the Stevens Institute of Technology for the use of AFM and Dr. 3812

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(34) Lapin, N. A.; Chabal, Y. J. J. Phys. Chem. B 2009, 113, 8776. (35) Chizhik, S. A.; Huang, Z.; Gorbunov, V. V.; Myshkin, N. K.; Tsukruk, V. V. Langmuir 1998, 14, 2605. (36) Sofia, S. J.; Premnath, V.; Merrill, E. W. Macromolecules 1998, 31, 5059. (37) Satulovsky, J.; Carignano, M. A.; Szleifer, I. Proc. Natl. Acad. Sci. U.S.A. 2000, 97, 9037. (38) Xue, C.; Yonet-Tanyeri, N.; Brouette, N.; Sferrazza, M.; Braun, P. V.; Leckband, D. E. Langmuir 2011, 27, 8810. (39) de Gennes, P. G. Macromolecules 1980, 13, 1069. (40) Daugaard, A. E.; Hvilsted, S.; Hansen, T. S.; Larsen, N. B. Macromolecules 2008, 41, 4321. (41) Bhushan, B.; Tokachichu, D. R.; Keener, M. T.; Lee, S. C. Acta Biomater. 2005, 1, 327. (42) Vesel, A.; Elersic, K.; Mozetic, M. Vacuum 2012, 86, 773. (43) Tsougeni, K.; Tserepi, A.; Constantoudis, V.; Gogolides, E.; Petrou, P. S.; Kakabakos, S. E. Langmuir 2010, 26, 13883.

Svetlana Sukhishvili for use of the contact angle instrument. We also thank Rob Planty at Rensselaer Polytechnic Institute (RPI) for XPS experiments.

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dx.doi.org/10.1021/la204773u | Langmuir 2012, 28, 3807−3813