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Fabrication of low-fouling, high-loading polymeric surfaces through pH-controlled RAFT

Low-fouling and high-loading surfaces are increasingly important for biosensing and blood purification technologies. Selective and efficient target binding from complex media can be achieved with poly(carboxybetaine) (pCB) surfaces that consist of a dense brush layer to resist non-specific protein adsorption and a sparse “mushroom” upper layer for high-density capture agent immobilization (i.e. high-loading). We developed pH-controlled surface-reversible addition–fragmentation chain-transfer (S-RAFT) polymerization to simplify fabrication of multi-modal, low-fouling and high-loading pCB surfaces without the need for quenching or re-initiation steps, toxic transition metals or light irradiation. Multi-modal polymer layers were produced through partial polymer termination by temporarily raising the pH to aminolyse a fraction of dormant chain transfer agents (CTAs); remaining polymer chains with intact CTAs continued uninterrupted extension to create the “mushroom” upper layer. The multi-modal pCB surfaces were low-fouling towards proteins (<6.7 ng cm−2), and macrophages. Compared to mono-modal brush surfaces, multi-modal pCB surfaces were high-loading with 5-fold greater capture agent immobilization (e.g. antibody) and 4-fold greater target binding (e.g. biotin-fluorescein).


Introduction
Bioactive polymeric surface coatings for low-fouling biointerfaces are being developed for biomolecule sensing in complex uids and blood purication. [1][2][3][4][5][6] To date, a variety of polymeric materials have been used to achieve low-fouling surfaces such as poly(ethylene glycol) (PEG), 7 uorinated polymers, 8 and zwitterionic polymers. 5 Poly(carboxybetaine) (pCB) has been identied as a zwitterionic polymer that both resists non-specic protein adsorption and is readily functionalized with biomolecules that act as capture agents for target molecule binding. Capture agents can be anchored to pCB's carboxylic acid groups while maintaining low-fouling properties towards proteins. 9 Furthermore, enzymes and antibodies conjugated to pCB maintain or increase in activity 10 and thermostability by restricting conformational transitions, 11 making it optimal for bioactive low-fouling surfaces.
Polymer surface coatings are produced by one of two methodologies: (1) "gra-to", where the end group of pre-synthesized polymer are covalently bonded to a surface; or (2) "gra-from", where polymerization is initiated from the surface. 12 Gra-to polymer densities are limited by the polymer chain's radius of gyration (R g ) leading to low density "pancake" or "mushroom" polymer conformations. 13 Gra-from yields polymer densities that exceed the R g imposed limit, and extends polymer chains in a brush conguration. 13 While gra-from's high polymer density improves fouling resistance, 14 biomolecule immobilization is limited to a monolayer on the brush surface; polymer side-chains are inaccessible due to tight polymer packing. 15 In contrast, gra-to's lower polymer density yields surfaces with greater loading potential but also increases fouling due to lower polymer surface coverage.
Bimodal architectures demonstrate high surface loading while remaining low-fouling, 9,15,16 combining the advantages of gra-from and gra-to. Bimodal polymer surfaces consist of a short dense polymer layer that resists protein fouling, and a sparse layer for greater biomolecule immobilization. Bimodal architectures have been synthesized previously via surface initiated-atom transfer radical polymerization (SI-ATRP) and surface initiated-photoiniferter mediated polymerization (SI-PIMP) 9,15,16 via two-step procedures that require termination and radical re-initiation steps, copper or surface exposure to light. The methods require two independent sequential radical polymerizations, where the rst polymerization is fully quenched before re-initiation of the synthesis of the second polymer layer.
To decrease synthetic complexity and expand accessibility, we developed pH-controlled S-RAFT for multimodal polymer architectures that avoids radical quenching, re-initiation and multiple CTA-immobilization steps (Fig. 1). The procedure terminates a subpopulation of CTAs during polymerization to establish the dense polymer layer, while the remaining chains with active CTAs continue to extend. Polymerization was conducted in the presence of a primary amine, butylamine, that is protonated and unreactive during S-RAFT polymerization at pH 4.5. To establish the rst layer, the pH is temporarily raised to 11 with NaOH, which deprotonates butylamine for partial CTA aminolysis. To create the sparse second layer, the pH is returned to 4.5 (by adding 12 M HCl) before complete CTA aminolysis. The bimodal pCB surfaces resisted non-specic protein adsorption (<6.7 ng cm À2 ) and decreased macrophage adhesion. Compared to brush only layers, the bimodal pCB layers increased antibody, capture agent, loading 5-fold and improved the capture of biotin-uorescein on avidin modied layers by 4-fold.
Carboxybetaine methacrylamide (CB) monomer was synthesized via a previously published method. 17 Briey, 23.25 g (136.5 mmol, 1 equiv.) of N- [3-(dimethylamino)propyl]methacrylamide was dissolved in 300 mL of dry acetonitrile under N 2 . tert-Butyl bromoacetate (30 g, 153.8 mmol, 1.1 equiv.) was added, and le to react overnight at 50 C. The reaction was cooled to room temperature and the product was precipitated with 500 mL of ether. The product was le to stand at 4 C overnight, and decanted. The white powder was collected, washed with 100 mL of ether, decanted, and dried under a stream of nitrogen, followed by incubation overnight under vacuum. The t-butyl protected intermediate was then deprotected by dissolving in neat TFA and incubating at room temperature for 2 h. The deprotected CB monomer was precipitated in ether, dried under a nitrogen stream, and freeze drying. 1   (s, 2H), 3.53 (m, 2H), 3.28 (t, J ¼ 6.42, 2H), 3.18 (s, 6H), 1.96 (m, 2H), 1.85 (s, 3H).
Fluorescent bevacizumab (bevacizumab-647) was synthesized by mixing 7.5 mL Alexa uor-647 NHS DMF solutions (10 mg mL À1 ; 0.075 mg, 0.06 mmol, 3 equiv.) with 100 mL of a bevacizumab solution (3 mg, 0.02 mmol, 1 equiv.) in PBS (pH 7.4) for 3 h in the dark. Bevacizumab-647 was puried by dialysis (MWCO 12-14 kDa) against PBS at 4 C in the dark. The nal bevacizumab-647 concentration and substitution ratio (dyes per antibody) was calculated from absorbance measurements taken with a Biotek Cytation 5 plate reader equipped with a Take3 micro-volume plate using extinction coefficient for Alexa uor-647 of 239 000 cm À1 M À1 and a correction factor of 0.03.

Modication of silica surfaces
2.2.1. Modication of silica surfaces and APTES deposition. Silicon wafers (100 mm, N-type, h100i, 1-10 ohm cm) were soaked in 1 : 1 HCl : methanol for 30 min, rinsed with Milli-Q water and dried under nitrogen. The wafers were then soaked in concentrated H 2 SO 4 for 30 min, rinsed with Milli-Q water and dried under a stream of nitrogen. Surfaces were then spin coated with a 0.1% v/v APTES in dry toluene (dried over 3Å molecular sieves), sonicated for 1 min in dry toluene, dried under a stream of nitrogen, and incubated for 1 h at 70 C.
2.2.2. Immobilization of RAFT chain transfer agent. 4-Cyano-4-(phenylcarbonothioylthio)pentanoic acid (14 mg, 0.05 mmol, 1 equiv.) was activated with DIC (39 mL, 0.25 mmol, 5 equiv.) and NHS (29 mg, 0.25 mmol, 5 equiv.) in DCM (1 mL; dried over 3Å molecular sieves) and stirred overnight at room temperature under nitrogen. APTES functionalized silica wafers were then added to the DCM solution, diluted to 5 mL, which was then kept under nitrogen for 60 h with gentle stirring. pentanoic acid (15.4 mg, 0.05 mmol, 1 equiv.), 4,4 0 -azobis(4cyanopentanoic acid) (7.7 mg, 0.27 mmol, 0.5 equiv.) and butylamine (50 mL, 0.5 mmol, 9 equiv.) were dissolved in 5 mL of 2 : 1 acetate buffer (0.1 M, pH 4.5) and 1,4-dioxane. The pH was adjusted to 4.5 with NaOH (8 M) and degassed via three freezepump-thaw cycles with nitrogen backlls and incubated at 70 C with gentle stirring. Two different bimodal polymerizations were performed: (1) aer 2 h at 70 C, an aliquot of polymerization solution was removed for gel permeation chromatography (GPC) characterization and degassed NaOH (8 M, 500 mL) was added to the reaction vessel to raise the pH and initiate CTA aminolysis. Aer 6 min at high pH, HCl (12 M, 335 mL) was added to end aminolysis and the polymerization was maintained for 22 h. And, (2) aer 4 h at 70 C, an aliquot of polymerization solution was removed for GPC characterization and degassed NaOH (8 M, 500 mL) was added to the reaction vessel to initiate CTA aminolysis. Aer 6 min at high pH, HCl (12 M, 335 mL) was then added to end aminolysis and polymerization was maintained for 20 h.
Trimodal pDMAPMA was synthesized similarly to the bimodal pDMAPMA polymerization described in condition 1 above, except degassed NaOH (8 M, 500 mL) was also added aer 9 h polymerizaton, and HCl (12 M, 335 mL) at 9 h and 6 min, with a total polymerization time of 45 h.
GPC chromatograms of bimodal pCB were deconvolved with Microso Excel assuming two peaks t as normal Gaussian distributions using the included generalized reduced gradient (GRG) nonlinear algorithm. The ratio of high to low molecular weight polymers in bimodal distributions were calculated as the relative ratio of the area of these Gaussian distributions. 2.4.2. X-ray photoelectron spectroscopy. Modied wafers were analyzed with a PHI Quantera II scanning X-ray photoelectron spectroscopy (XPS) microprobe. A take-off angle of 45 was used for all samples, pass energy and step size were 224 eV and 0.8 eV for survey scans and 55 eV and 0.1 eV for high resolution scans, which were used to determine elemental composition.
2.4.3. Water contact angle. Static water contact angle (WCA) measurements and images were acquired with an OCA 20 (Future Digital Scientic) contact angle measurement system and calculated with the SCA 20 soware module. 3 mL droplets of Milli-Q water were deposited on APTES functionalized, CTA functionalized, and 1 and 2 layer pCB surfaces. Surface uorescence microscopy imaging was performed with a Biotek Cytation microscope, equipped with a Cy5 lter set. Intensity quantication was performed with the ImageJ Mean Gray Value tool. Surface uorescence values were converted to ng cm À2 by comparison with a calibration curve of dropcast bevacizumab-647 of known total mass onto wafers and imaged under the same conditions as experimental samples to produce a calibration curve.
2.5.2. Avidin immobilization and biotin uorescein capture. Avidin was immobilized onto 1 and 2 layer pCB using the same procedure as bevacizumab-647. Following immobilization, surfaces were rinsed 6 times with 0.05% TWEEN 20 in PBS (PBS-T). Avidin functionalized wafers and pCB coated wafers without avidin were incubated in 100 mL of 4 mg mL À1 biotin-uorescein in PBS for 2 h in the dark under agitation (orbital shaker at 100 RPM). The biotin-uorescein solution was removed and surfaces were rinsed 6 times with PBS-T. Fluorescence of each well was measured with a Biotek Cytation plate reader and surface uorescence was imaged by uorescent microscopy. Surface uorescence values were converted to ng cm À2 by comparison with a calibration curve of dropcast biotin-uorescein of known total mass onto wafers and imaged under the same conditions as experimental samples.
2.6. Characterization of protein and macrophage fouling on pCB coated surfaces 2.6.1. Non-specic protein fouling. Total non-specic protein adsorption was measured by a bicinchoninic acid (BCA) assay of collected protein, as previously reported. [18][19][20] Pristine silicon, 1 and 2 layer pCB, and pDMAPMA surfaces were incubated with 100% aged CBS at 37 C for 1 h. Following incubation, the wafers were rinsed with deionized (DI) water and incubated in PBS for 15 min on an orbital shaker table at 100 RPM. The wafers were then removed from the PBS solution and dried under a stream of nitrogen. To extract adsorbed protein from the dry wafers, 120 mL of 8% SDS in deionized (DI) water was added to cover the surface, the wafers were then incubated in a humidity chamber overnight at 37 C. The SDS solution was removed from the wafers and incubated for 2 h at 37 C with BCA assay kit reagents as per manufacturer protocol. 21 Finally, absorbance was read at 562 nm and compared against a BSA calibration curve to determine the total adsorbed protein content. The total amount of protein in the SDS solution was then determined using the BCA protein quantication assay with a calibration curve. The calibration curve was prepared using BSA calibrants in the SDS solution used to remove adsorbed protein from wafers. The lowest detectable concentration in the BSA calibration curve was 0.3 mg mL À1 . The SDS solutions used for BSA extraction were 120 mL and the wafers were 5.4 cm 2 . Therefore, the limit of detection (LOD) is equal to: 2.6.2. Non-specic macrophage adhesion. pCB coated silicon wafers were placed into a 96 well plate. Wells were soaked in 70% ethanol for 1 h to sterilize the surfaces. The ethanol was removed and each wafer was rinsed three times with water, and incubated for 24 h with 200 mL of FBS at 37 C. FBS was removed and 200 mL of RAW 264.7 macrophages (50 000 cells per mL) in DMEM supplemented with 10% FBS were added to each well and cultured for 2 d at 37 C and 5% CO 2 . Cell media was then removed, and each well was rinsed three times with 37 C PBS. Cells were stained with calcein AM and HOESCHT according to manufacturer protocols. Silicon wafers were imaged with a Biotek Cytation uorescent microscope equipped with DAPI and GFP lter cubes, and cells with colocalized stains were counted.

Statistical methods
All graphed data represents mean AE standard deviation calculated from triplicate measurements, unless otherwise stated. All statistical analysis was performed using GraphPad Prism 8. P < 0.05 is indicated by *, P < 0.01 by **, and P < 0.001 by ***. To demonstrate the production of multimodal polymer distributions using pH-controlled RAFT, we rst conducted and characterized the solution polymerization of pDMAPMA and pCB, using a thiocarbonylthio containing CTA end group at the polymer's living end. Controlled partial degradation of the thiocarbonylthio group, which prevents further monomer addition, led to bi-or multi-modal polymer M w distributions. Here, This journal is © The Royal Society of Chemistry 2020 we reacted a fraction of polymer thiocarbonylthio end groups with an amine nucleophile, butylamine, to yield terminal thiols and a thioamide side product (Scheme S1 †).

Results and discussion
Partial termination of CTAs with amino nucleophiles allows for one-pot synthesis of bimodal polymer M w distributions by temporarily raising the pH; the amine is protonated and unreactive at pH 4.5, and unprotonated and reactive at pH 11. Butylamine was chosen as the nucleophile because it is soluble under pCB's aqueous polymerization conditions and has a pK a of $10.5, 22 which minimizes aminolysis during pH 4.5 RAFT polymerization. 23,24 A 10-fold molar excess of butylamine compared to CTA agent during polymerization ensured efficient aminolysis at high pH. The elapsed polymerization time prior to pH raising determined the M w of the low M w polymer population.
Two different bimodal pDMAPMAs were synthesized with partial aminolysis at 2 or 4 h aer initiation; polymers were characterized by GPC ( Fig. 2A) and NMR (Fig. S1B †). The low M w polymer population increased from 8.0 to 9.2 kDa as polymerization time prior to aminolysis increased from 2 to 4 h, minimal differences in Đ were observed (1.03 versus 1.05; Table  1). We also conrmed that pDMAPMA's tertiary amine does not result in CTA aminolysis at pH 4.5 (Fig. S2 †). The ability to synthesize multimodal distributions was further tested by synthesizing a trimodal distribution of pDMAPMA with partial aminolysis at 2 and 9 h, and a total polymerization time of 45 h. This resulted in the production of three distinct M w populations of 7 (Đ ¼ 1.02), 70, and 391 kDa, respectively; the dispersity of the two high M w populations could not be accurately calculated due to elution near the exclusion limit of the column (Fig. 2B).
To determine the duration of aminolysis required for pCB bimodal distributions, pCB was synthesized with pH 11 intervals of 5, 30 or 60 min. The 5 min interval yielded a clear bimodal distribution according to GPC analysis, with two M w populations of 6.5 (Đ ¼ 1.02) and 10.5 kDa (Đ ¼ 1.11) (Table 1, Fig. 3A); the bimodal solution was further characterized by NMR (Fig. S1A †). The 30 and 60 min intervals most likely resulted in maximum aminolysis of the CTAs and uncontrolled polymerization of propagating chains, which resulted in bimodal distributions dominated by the low M w population (Fig. 3). To determine the  relative distribution of low and high M w populations, the GPC data was deconvolved assuming Gaussian distribution to yield a 0.43 : 0.57 ratio of low to high M w populations for the 5 min pH 11 time interval. Deconvolution of GPC traces of conditions with 30 and 60 min at pH 11 yielded greater proportions of the low M w population, as expected. To ensure aminolysis is required for bimodal polymerization, the pCB synthesis protocol was repeated in the absence of butylamine; no clear bimodal pCB distributions were observed in the absence of butyl amine (Fig. S3 †). Therefore, bimodal populations of pCB can be synthesized using pH-controlled RAFT with butylamine and a pH 11 interval of 5 min. Because the monomer must be stable under basic conditions for pH controlled aminolysis, CB with a 1 carbon spacer length (CSL) between charges was chosen; CBs with a CSL of 2 have been reported to undergo elimination reactions under basic conditions. 25 We conrmed that no pCB structure modications occurred upon exposure to high pH aminolysis conditions by NMR (Fig. S1 †).

Gra-from pCB synthesis on silicon wafer surfaces was conrmed by XPS
Using XPS, we rst conrmed that the silicon wafers were activated with CTA agent and amenable to gra-from polymerization of pCB. Silicon wafers were spin coated with APTES and reacted with CTA's carboxylic acid using DIC to form an unreactive amide bond towards aminolysis. XPS of the wafers conrmed the presence of nitrogen and sulfur aer immobilization of APTES and CTA, respectively, which indicates successful surface functionalization. The wafers were then immersed in a pCB polymerization solution that was then heated to initiate concurrent solution and surface pCB polymerization. Aer polymerization, the pCB modied wafers had greater carbon content compared to CTA modied surfaces according to XPS (Table 2). Therefore, the CTA modied wafers are amenable to pCB gra-from polymerization.

Gra-from synthesis of bimodal pCB through pHcontrolled S-RAFT
The synthesis of bimodal pCB surfaces using pH-controlled S-RAFT was conrmed through WCA, spectral ellipsometry, and  protein loading. Bimodal pCB layers were successfully polymerized with a 5 min pH 11 interval aer 1 h of polymerization, as described in Section 3.1. As controls, 1 layer pCB wafers were prepared with a total polymerization time of 1 h at pH 4.5.
To follow each surface modication step, the WCA of the wafers was measured aer APTES, CTA, 1 layer pCB and 2 layer pCB synthetic steps. APTES modied surfaces demonstrated a WCA of $40 that is characteristic of an aminosilane monolayer on silicon surfaces, 26 where the amines are orientated away from the surface with all three ethoxy groups reacted. CTA functionalized surfaces demonstrated WCAs ($55 ) typical of CTA, 4-cyano-4-(phenylcarbonothioylthio)pentanoic acid, monolayers. 27 The WCA of pCB modied surfaces is dependent on the degree of polymerization; 28 therefore, 2 layer (bimodal) pCB surfaces with a 24 h polymerization time had a lower WCA than 1 layer pCB surfaces with a 1 h polymerization time. The WCA of the bimodal architecture ($10 ) is hydrophilic and equivalent to the those previously reported for thick pCB layers. 29 The high variability in WCA for 1 layer pCB modied wafers (standard deviation of 10, Fig. 4A) is characteristic of thin coatings due to the sensitivity of WCA measurements to slight lm thickness variations. 28 To demonstrate that polymerization time (i.e. degree of polymerization) inuences WCA measurements, the WCA of unimodal and bimodal pCB surfaces produced using a 24 h polymerization timed had similar WCAs ( Fig. 4A and S4 †).
The greater thickness of 2 layer pCB architectures compared to 1 layer pCB was conrmed by spectral ellipsometry of dry pCB lms. The average thickness of the 2 layer coating was $2.4 times greater than the 1 layer coating (Fig. 4B), due to the formation of the low-density second layer. The layer thickness increased at each modication step from the native oxide layer on the silicon wafer, APTES, CTA, 1 layer pCB and 2 layer pCB (Fig. 4B), indicating a successful modi-cation procedure.
The 2 layer pCB architecture was also conrmed to have greater antibody loading potential through the immobilization of a uorescent antibody, bevacizumab modied with Alexa 647 (Fig. 5). Bimodal pCB layers have previously been shown to increase the surface capacity of immobilized antibodies when compared to monomodal brush pCB layers, where thicker low density layers lead to greater loading levels. 16 High-density monomodal polymer layers demonstrate lower degrees of protein immobilization because high-density of polymer chains prevent biomacromolecule diffusion. 30 Using EDC/NHS chemistry to activate pCB carboxylic acids, Alexa 647 modied bevacizumab was immobilized onto monomodal brush and bimodal pCB surfaces. Wafers were then characterized by surface uorescence, which indicated $5-fold greater bevacizumab immobilization on bimodal surfaces. Therefore, bimodal pCB surfaces produced through pH controlled S-RAFT retained the advantageous property of greater capture agent immobilization. Fig. 4 Surface characterization of functionalized silicon wafers by WCA and ellipsometry. (A) Static WCA and representative photographs of 3 mL water droplets on wafers functionalized with aminosilane, CTA, and 1 and 2 layer pCB coatings. Hydrophobic CTA increased the WCA contact, whereas increasing pCB content decreased the WCA. The WCA of 1 layer pCB was greater than 2 layer pCB due to differences in layer thickness. (B) Modeled layer thickness of native oxide, APTES, CTA, and 1 and 2 layer pCB coatings from spectral ellipsometric measurements. Statistics performed by one-way ANOVA with Bonferroni's multiple comparison test (mean AE standard deviation p < 0.05 (*), and p < 0.001 by (***)).

Fig. 5
Immobilization of bevacizumab confirmed the greater loading capacity of bimodal pCB wafers. Fluorescently labelled bevacizumab was immobilized using EDC/NHS chemistry on 1 (brush) and 2 layer (bimodal) pCB wafers. After quantification by fluorescent microscopy, the 2 layer pCB architecture was confirmed to have a higher loading capacity than 1 layer pCB. P < 0.01 (**), by Student's t-test, mean AE standard deviation.

Avidin modied bimodal pCB surfaces enhanced biotin capture
Bimodal pCB layers functionalized with avidin captured approximately 4 times more biotin-uorescein than avidin modied brush pCB surfaces (Fig. 6). The improved capturing capacity is similar to previously reported bimodal polymer surfaces which increased capacity by 1.8 to 3.1 times. 9,15 Avidin was rst immobilized on the pCB surfaces using EDC/NHS chemistry. The surfaces were then incubated in biotin-uorescein solutions for 2 h, then extensively washed to remove unbound biotin. Fluorescence microscopy was performed on biotin-uorescein exposed 1 and 2 pCB layer surfaces with and without immobilized avidin (Fig. 6B). The bimodal 2 layer pCB surfaces with avidin had greater total uorescence intensity, which indicates a greater amount of captured biotin-uorescein compared to 1 layer pCB surfaces. Quantitative surface uorescence measurements indicated $4-fold increase in biotin-uorescein capture on bimodal (2 layer) architectures when compared to brush (1 layer) surfaces (Fig. 6C); surfaces without avidin were used as background controls. Therefore, bimodal architectures improved both capture agent immobilization and target molecule binding.
3.5. pCB surfaces prepared by pH-controlled RAFT are lowfouling 1 and 2 layer pCB architectures had similar low-fouling properties due to the pCB brush architecture present on both surfaces. pCB coatings synthesized by the gra-from method are employed to generate low-fouling surfaces towards proteins, which has been dened as below 5 ng cm À2 ; 31 whereas, surfaces prepared by gra-to methods exposed to undiluted serum non-specically adsorb proteins at 11 ng cm À2 . 32 In this study, 1 and 2 layer pCB architectures were exposed to 100% aged CBS to evaluate protein fouling in complex media relevant for biological applications; aged serum has been shown to foul surfaces to a greater degree than fresh serum. 33 pCB surfaces were compared to positive fouling controls of cationic pDMAPMA modied surfaces and anionic plasma cleaned SiO 2 wafers. All surfaces were incubated in CBS for 1 h and washed extensively in PBS prior to adsorbed protein quantication using an SDS extraction method and the BCA protein quantication assay, as previously described. 20 Adsorbed protein on the 1 and 2 layer pCB surfaces was below the assay detection limit of 6.7 ng cm À2 , while the pDMAPMA control surface had $130 ng of adsorbed protein per cm À2 (Fig. 7A). The calculated values of adsorbed Fig. 6 Bimodal, two layer, pCB wafers modified with avidin have greater biotin capturing efficiency. (A) Schematic for the modification of pCB surfaces with avidin and subsequent biotin-fluorescein capture. (B) Representative fluorescent micrographs (4Â magnification) and micrograph fluorescent intensity from 1 and 2 layer pCB surfaces with and without avidin that were exposed to biotin-fluorescein. (C) Surface fluorescence quantification of biotin-fluorescein bound to avidin modified surfaces (n ¼ 5 for 1 layer pCB, n ¼ 6 for 2 layer pCB, mean AE standard deviation, p < 0.05 (*), Student's t-test). Background signal was removed by measuring fluorescence of avidin-free surfaces after exposure to biotinfluorescein.
protein assumed complete extraction with SDS, which has been shown to not hold for all protein interactions, 34 and thus represent lower bounds. Nonetheless, 1 and 2 layer pCB coatings were demonstrated to improve protein fouling resistance. The relative decrease in detected absorbed protein between control surfaces and pCB coatings is similar to previously reported values using the SDS BCA assay ($20Â). 35,36 Therefore, bimodal pCB surfaces synthesized by pH-controlled S-RAFT resist protein fouling to a similar extent as previously reported bimodal architectures synthesized by SI-ATRP and SI-PIMP.
Surface fouling was further characterized by quantifying non-specic macrophage adhesion to serum exposed surfaces. Both 1 and 2 layer pCB coatings reduced macrophage adhesion by >90% when compared to tissue culture plastic (TCP) and plasma cleaned wafers (SiO 2 ; Fig. 7B and C). To ensure maximum macrophage binding, all surfaces were pre-incubated in FBS for 24 h prior to cell seeding (10 000 macrophages per surface). Aer seeding, surfaces were cultured for an additional 2 d to promote maximum macrophage adhesion to identify potential difference in nonspecic macrophage adhesion. Surfaces were then gently washed with PBS and cells were stained with HOESCHT and calcein AM for counting. Both 1 and 2 layer pCB decreased macrophage adhesion to a similar extent ( Fig. 7B and C), indicating both surfaces had similar lowfouling properties towards cells, as expected.

Further discussion
pH-Controlled bimodal S-RAFT polymerization simplies the synthesis of bimodal pCB layers and avoids the use of potentially toxic materials (e.g. Cu), increasing the accessibility of bimodal pCB surfaces for incorporation into sensors and medical devices. Bimodal architectures were achieved by the temporal control of pH (through NaOH and HCl additions) to temporarily deprotonate butylamine for partial CTA aminolysis. The method can be extended to other CTAs due to the standard thiocarbonylthio bond found in many RAFT CTAs, 37 enabling polymerization of monomers that may not be compatible with 4-cyano-4-(phenylcarbonothioylthio)pentanoic acid. This technique can further be extended to other substrate materials such as gold, an important substrate for biosensors, as only the initial amine functionalization chemistry must be adapted; S-RAFT has previously been performed on gold surfaces. 38 The development of high loading, low-fouling surfaces is particularly important for biomedical sensors and blood Fig. 7 One and two layer pCB surfaces were equally low-fouling when exposed to serum. (A) Non-specific protein adsorption onto pristine silicon wafers and wafers functionalized with single layers of pDMAPMA and pCB, and bimodal pCB when incubated in CBS for 24 h; cationic pDMAPMA and anionic SiO 2 surfaces were included as a positive fouling control. No adsorbed protein was detected on 1 or 2 layer pCB (detection limit of the assay was 6.7 ng cm À2 ); protein adsorption was therefore comparable to non-fouling surfaces. 31 . (B) After 48 h of culturing, relative amounts of adhered macrophages on surfaces pre-exposed to serum for 24 h was quantified. Cell counts were determined from fluorescent micrographs; numbers were normalized to the tissue culture plastic (TCP) control. (C) Representative fluorescent micrographs of adhered macrophages stained with HOESCHT and calcein AM, scale bars ¼ 200 mm. purication devices. The ability to rapidly detect cytokines in blood or tissue extracts is increasingly important due to the development of immune modulating drugs such as cancer immunotherapies, where cytokine levels are carefully tracked. 39 Recently, the development of polymeric surfaces for blood cytokine 40 and lipopolysaccharide 41 (LPS) ltration has shown promise for the treatment of sepsis; cytokine ltration could also have applications in cardio-renal syndromes. 42 Blood ltration with low-fouling bioactive surfaces can selectively alter blood biochemistry with temporal control by administering inline lters to patients. 43

Conclusions
The development of low-fouling surfaces that selectively capture molecules from biological uids is crucial in the development of biosensors and blood ltration devices. We developed pHcontrolled S-RAFT to simplify the synthesis of bimodal pCB layers that increase capturing efficiency while retaining lowfouling properties. Compared to monomodal pCB architectures, bimodal surfaces improved capture agent, antibody, immobilization and target molecule, biotin-uorescein, capture by 4 to 5-fold. pH-Controlled S-RAFT only requires a temporary increase in pH for CTA aminolysis during surface preparation, which will help towards increasing the accessibility of bimodal architectures for biomedical applications.

Conflicts of interest
There are no conicts to declare.