We designed non-textured, all-solid, slippery hydrophilic (SLIC) surfaces by covalently grafting polyethylene glycol brushes to smooth solid substrates above a critical grafting density, which occurs when tethered brush size is equal to intertether distance. Our SLIC surfaces demonstrated exceptional performance in condensation and fouling resistance compared with non-slippery hydrophilic and slippery hydrophobic surfaces. We firmly believe that SLIC surfaces constitute an emerging class of surfaces with the potential to benefit multiple technological landscapes ranging from thermofluidics to biofluidics.

Slippery surfaces (i.e., surfaces enabling high mobility of liquids) have received significant attention due to their wide range of applications in self-cleaning, drag reduction, enhanced condensation, biomedical implants, etc. [1][2][3][4][5][6] In the past 2 decades, most slippery surfaces have been fabricated using two approaches, both of which rely on principles of fluid-film lubrication. 7 The first approach employs super-repellent surfaces, [8][9][10][11][12] which use texture to trap pockets of air at the solid surface. The trapped air acts as a gaseous lubricant, reducing the solid-liquid interfacial area and inducing slip at the liquid-air interface. 13,14 The second approach employs lubricant-infused surfaces, 15,16 which use texture to trap pockets of a liquid, immiscible with the contacting liquids, at the solid surface. The trapped immiscible liquid acts as a liquid lubricant, reducing the solid-liquid interfacial area and inducing slip at the liquid-liquid interface. 17 Despite the appeal of super-repellent surfaces and lubricant-infused surfaces, they lose slipperiness due to damage of texture, depletion of air via dissolution or external pressure, or depletion of lubricant via evaporation or repeated use. [18][19][20][21][22] To circumvent these issues, non-textured, all-solid slippery surfaces have been developed by covalently grafting brushes to substrates (i.e., tethering oligomers or polymers to surfaces). 23,24 Since these surfaces are non-textured and all-solid, concerns related to damage of texture and depletion of fluid are greatly mitigated. Nearly all non-textured, all-solid, slippery surfaces are hydrophobic. [23][24][25][26][27] There are very few reports of non-textured, all-solid, slippery

Water droplets tend to move and slide easily on non-stick coatings. Such slipperiness is intuitive because almost all non-stick coatings have hydrophobic (i.e., water-repellent) surfaces, which do not stick much to water. On the contrary, water spreads and sticks easily to hydrophilic (i.e., waterloving) surfaces. So, hydrophilic surfaces, which allow water droplets to move and slide easily, are counter-intuitive and rare. In this work, we divulge the design principles for making such slippery hydrophilic (SLIC) surfaces. Such SLIC surfaces constitute an emerging class of surfaces with significant potential to benefit multiple technological landscapes ranging from thermofluidics to biofluidics. hydrophilic surfaces, [28][29][30][31] and no studies have systematically investigated or elucidated the design of such surfaces. In this work, we demonstrate that non-textured, all-solid, slippery hydrophilic surfaces can be designed by covalently binding high-surface-energy brushes to smooth substrates, only at sufficiently high grafting densities, resulting in both physical and chemical homogeneity. Furthermore, we postulate a plateau in slipperiness above a critical grafting density, which occurs when the tethered brush size (i.e., twice the Flory radius 32 ) is equal to the inter-tether distance. This plateau in slipperiness occurs because, at or above the critical grafting density, water molecules tend to move past the tethered brushes primarily via inplane motion, with negligible penetration between the brushes, resulting in maximum and constant slipperiness. Building on this understanding, we fabricated slippery hydrophilic (SLIC) surfaces by covalently binding polyethylene glycol (PEG) brushes to smooth substrates with a rapid (<10 min) grafting-to approach. Our SLIC surfaces demonstrate exceptional performance in condensation and fouling resistance compared with conventional (i.e., non-slippery) hydrophilic surfaces and slippery hydrophobic surfaces. Based on these results, non-textured, all-solid, slippery hydrophilic surfaces (e.g., SLIC surfaces) constitute an emerging class of surfaces with the potential to benefit multiple technological landscapes ranging from thermofluidics to biofluidics.

Hydrophilic surfaces can be fabricated by covalently attaching high-surface-energy oligomeric brushes (e.g., PEG brushes) to a solid surface. 33 Such hydrophilic surfaces can display slipperiness when contact angle hysteresis Dq (i.e., the difference between advancing contact angle q adv and receding contact angle qrec ) is low. Low-contact-angle hysteresis can be achieved when a surface has low physical and chemical inhomogeneities. 34,35 Non-textured (i.e., smooth) surfaces with very low surface roughness display low physical inhomogeneity. High grafting density s leads to low chemical inhomogeneity. When hydrophilic P EG brushes are grafted to smooth surfaces, slipperiness increases with increasing grafting density. This is because, at lower grafting densities, water molecules tend to penetrate the space between the PEG brushes (see Figure 1A), resulting in pinning and higher hindrance to lateral mobility of water droplets, which in turn leads to lower slipperiness (i.e., higher Dq). As grafting density increases, the penetration of water molecules between the PEG brushes decreases, resulting in lower pinning and lower hindrance to lateral mobility of water droplets, which in turn leads to increasing slipperiness (i.e., decreasing Dq). We postulate that at or above a critical grafting density (i.e., s R scrit ), where the brush size (i.e., twice the Flory radius, 2RF ) is equal to the inter-tether distance, slipperiness is maximum (i.e., Dq is minimum) and constant. This is because, at or above the critical grafting density, water molecules tend to move past the oligomeric brushes primarily via in-plane motion with negligible penetration (see Figure 1B), possibly due to an ice-like hydration layer. [36][37][38] These physical insights can be conveniently expressed in terms of the non-dimensional slipperiness factor s = s= scrit . When s < 1, it signifies a regime where slipperiness increases with increasing graft-ing density, and when s R 1, it signifies a regime of maximum and constant slipper-iness. So, s > 1 is a prudent criterion for designing non-textured, all-solid, slippery hydrophilic surfaces.

Building on this understanding, in this work, we fabricated SLIC surfaces by covalently grafting PEG brushes (molecular weight, M z 330 Da) to silicon wafers using a rapid grafting-to approach (see Figure 1C). We chose silicon wafers as substrates to minimize physical inhomogeneities (i.e., surface roughness, Rrms). We chose PEG p ffi ffi ffi functional groups to render high solid surface energy, which makes the surfaces hydrophilic. We used a simple and rapid liquid phase silanization 39 (see experimental procedures) to covalently graft PEG brushes (see XPS spectra, Figure 2A) to hydroxylated silicon wafers, while maintaining low surface roughness (Rrms < 1 nm; see Figure 2B). To understand the influence of grafting density on wettability and slipperiness, we systematically tuned the non-dimensional slipperiness factor s by tailoring the silanization time t sil (see Note S1 and Table S1). At different t sil , we estimated the grafting density (in chains nm 2 ), s = ðhrNA x 10 21 Þ=M; here, h is the thickness of the P EG layer (in nm) estimated from ellipsometry (see experimental procedures), r is the density of PEG (in g cm 3 ), NA is Avogadro's number, and M is the number averaged molecular weight of PEG (in Da). We estimated the critical grafting density scrit = 1=ð2 3 R 2 Þ, assuming hexagonal packing with a grafting density s. At t sil = 0 min, water droplets completely spread (static contact angle, q = 0) and did not slide on hydroxylated silicon wafers. As tsil increased, the dynamic contact angles (i.e., q adv and qrec ) of water increased (see Figure 2C), indicating increasing covalent grafting of PEG brushes, resulting in increasing s. In addition, as t sil increased, Dq decreased (see Figure 2D), indicating a decrease in inhomogeneities on the sur-face as s / 1. At t sil = 7 min, we estimated that s > 1 for P EG brushes, indicating the onset of maximum and constant slipperiness regime with q adv = 39, qrec = 36, and Dq = 3. For t sil R 7 min, we obtained SLIC surfaces (see Figure 2E and Video S1) with nearly constant contact angles and contact angle hysteresis, confirming that s > 1 signifies a regime of maximum and constant slipperiness. In a similar manner, we also fabricated SLIC surfaces using longer PEG chains (M z 462 Da; see Note S2 and Table S2) at s > 1 reaffirming the criterion for designing non-textured, all-solid, slippery hydrophilic surfaces.

In addition to elucidating the underlying design principles and fabricating SLIC surfaces, we evaluated the performance of our SLIC surfaces in relevant thermofluidic and biofluidic applications. Specifically, our SLIC surfaces are ideal for enhancing condensation heat transfer of aqueous liquids because they simultaneously offer low static contact angle q and low sliding angle u (i.e., minimum tilt angle for a droplet to slide on a surface) for aqueous liquids. Low contact angle q promotes high nucleation rates and low conduction resistance, both of which enhance condensation heat transfer. 31,40,41 Low sliding angle u (due to low contact angle hysteresis Dq) promotes rapid and efficient removal of the condensate droplets, thereby enabling sustained dropwise condensation with significantly higher heat transfer coefficient compared with film-wise condensation. 31,40,41 To demonstrate such enhanced condensation heat transfer, we qualitatively compared the condensation of water (see experimental procedures) on our SLIC surfaces (q z 37, u z 3 for 20 mL water droplets) with condensation on non-slippery hydrophilic surfaces (having low contact angle, but higher sliding angle) and slippery hydrophobic surfaces (having low sliding angle, but higher contact angle). We used untreated silicon wafers as non-slippery hydrophilic surfaces (q z 45, u z 20 for 20 mL water drop-lets) and 1,3-dichlorotetramethyldisiloxane-treated silicon wafers as slippery hydrophobic surfaces (q z 103, q adv z 104, qrec z 101, u z 3 for 20 mL water droplets; see experimental procedures). We exposed vertically mounted surfaces at ambient temperature ( z 2 0 C ) to steam at approximately 100C (see Figures 3A-3C and Video S2). The non-slippery hydrophilic surface and our SLIC surface displayed faster droplet coalescence and growth due to their higher nucleation rate compared with the hydrophobic slippery surface. Condensate droplets attained a critical size ( z 1 . 4 mm; see Figure 3C) and slid down our SLIC surface at a faster rate compared with the hydrophobic slippery surface and the non-slippery hydrophilic surface. This rapid droplet shedding indicates that our SLIC surfaces not only facilitate high nucleation rates due to their hydrophilicity but also facilitate rapid condensate removal due to their high slipperiness (i.e., low u). After removal of a condensate droplet, the nucleation, coalescence, growth, and removal of additional condensate droplets continued on our SLIC surface, indicating the potential for sustained dropwise condensation and enhanced condensation heat transfer. In practice, condensation can occur at different relative humidities for a wide range of aqueous liquids. To ensure that our SLIC surfaces retain their hydrophilicity and slipperiness, we characterized their contact angles and sliding angles for water at different relative humidities and for aqueous liquids (solutions of sodium dodecyl sulfate in water) with a wide range of surface tensions (g lv = 40 to 72 mN m 1 ; see experimental procedures). Our SLIC surfaces displayed negligible change in static contact angles and sliding angles of water at relative humidities ranging from 10% to 100% (see Figure 3D). As the surface tension of aqueous liquids decreased from 72 to 40 mN m 1 , our SLIC surfaces displayed an expected decrease in static contact angles and sliding angles (see Figure 3E, Note S3 and Table S3). These results demonstrate the retention of hydrophilicity and slipperiness of our SLIC surfaces and their poten-tial for condensation applications. In addition, our SLIC surfaces also demonstrated the retention of hydrophilicity and slipperiness even after exposure to air for 20 days, immersion in water for 20 days, exposure to steam (at 100C and 1 atm) for a cumulative time of 48 h and 100,000 water droplets sliding past the surface (see Note S4).

In addition to thermofluidic applications, we also evaluated the performance of our SLIC surfaces in biofluidic applications. Specifically, our SLIC surfaces are ideal for fouling-resistant lab-on-a-chip miniature platforms because they simultaneously offer hydrophilicity and slipperiness for aqueous liquids. Hydrophilicity imparted by the PEG brushes allows improved resistance to fouling by proteins because of the hydration layer and the steric hindrance. [42][43][44][45] To demonstrate the fouling resistance of our SLIC surfaces, we compared them with that of non-slippery hydrophilic surfaces and slippery hydrophobic surfaces. We used untreated glass coverslips as non-slippery hydrophilic surfaces (q z 53, u z 47 for 20 mL water droplets), 1,3-dichlorotetramethyldisiloxane-treated glass coverslips as slippery hydrophobic surfaces (q z 103, u z 4 for 20 mL water droplets), and glass cover-slips with covalently grafted PEG brushes as our SLIC surfaces (q z 37, u z 5 for 20 mL water droplets). We used glass coverslips because they are smooth (allowing low physical inhomogeneity) and transparent (allowing light microscopy). We exposed the surfaces to fibrinogen solution and studied protein adsorption in situ using total internal reflection fluorescence (TIRF) microscopy (see experimental procedures). Fibrinogen adsorbed within 1 min on both the non-slippery hydrophilic surfaces and slippery hydrophobic surfaces (see Figures 4A and4B, and Video S3). In contrast, our SLIC surfaces displayed exceptional fouling resistance by preventing adsorption of fibrinogen even at 1,800 min (see Figure 4C and Video S3). 46,47 The exceptional fouling resistance is possibly due to an ice-like hydration layer, 37,38,43 but a comprehensive study is needed to develop a mecha-nistic understanding. In addition to the outstanding fouling resistance, the slipper-iness of our SLIC surfaces allows aqueous droplets to slide easily (see Figure 4D), allowing manipulation of aqueous droplets. To demonstrate the utility of slipper-iness in labon-a-chip miniature platforms, we fabricated SLIC surfaces patterned with an array of non-slippery domains (see experimental procedures). When an aqueous droplet (e.g., protein solution) slides past such patterned SLIC surfaces, tiny volumes of the droplet are trapped in the non-slippery domains due to higher adhesion (see Figure 4D and Video S4). In this manner, on patterned SLIC sur-faces, biological analytes (e.g., proteins, nucleic acids, cells, microorganisms, etc.) can be trapped in targeted domains while ensuring that the remaining areas are free of fouling (see Figure 4D). Furthermore, while aqueous droplets slide easily, oil droplets adhere to our SLIC surfaces. When a compound droplet, con-sisting of water and oil, is placed on a tilted SLIC surface, water displaces oil and slides away, while the oil remains adhered (see Figures 4E-4G, Video S5, and Note S5), indicating the potential for biofluidic separations.

In this work, we elucidated the design of non-textured, all-solid, slippery hydrophilic surfaces. We demonstrated two distinct regimes of slipperiness: s < 1, where slipperiness increases with increasing grafting density, and s R 1, where slipperiness is maximum and constant. We postulate that s > 1 is a prudent criterion for the design of non-textured, all-solid, slippery hydrophilic surfaces. Building on this design criterion, we fabricated SLIC surfaces by covalently grafting PEG brushes to silicon wafers and glass coverslips. We demonstrated the exceptional performance of our SLIC surfaces in condensation and fouling resistance compared with non-slippery hydrophilic and slippery hydrophobic surfaces. While a comprehensive study (on processing-structure relationships with a wide variety of chemistries and substrates and the role of ice-like hydration layers) is needed to develop a thorough understanding of this emerging class of surfaces, our results indicate that SLIC surfaces have the potential to positively transform the technological landscape of thermofluidic and biofluidic devices.

Resource availability Lead contact Further information and requests for resources and materials should be directed to and will be fulfilled by the lead contact, Arun K. Kota (akota2@ncsu.edu).

This study did not generate new unique reagents.

The data generated during this study are available within the article and its supplemental information files and also are available from the corresponding authors upon reasonable request.

Fabrication of hydrophilic and hydrophobic surfaces Silicon wafers (<100> orientation; University Wafers) and glass cover slips (Fisher) were cleaned by rinsing thoroughly with acetone (Fisher) and DI water and then dried with nitrogen. The cleaned substrates were exposed to oxygen plasma (Plasma Etch PE-25) for 10 min for hydroxylation. To prepare hydrophilic surfaces, the hydroxylated samples were immersed in a solution consisting of 2 mL of 2methoxy polyethyleneoxy (6-9) propyl trimethoxysilane (Gelest) and 12 mL hydrochloric acid (Fisher) in 45 mL of anhydrous toluene (Fisher) for the desired time at room temperature. We chose low-molecular-weight PEG silane to reduce steric hindrance and obtain higher grafting densities. To prepare hydrophobic surfaces, the hydroxylated samples were exposed to the vapors of 150 mL of 1,3-dichlorotetramethyldisiloxane (Gelest) in an enclosed chamber for 15 min at room temperature. Finally, the silanized samples were cleaned by rinsing thoroughly with anhydrous toluene, DI water, and ethanol (Fisher) sequentially and then drying with nitrogen.

Star-shaped non-slippery domains were fabricated on SLIC surfaces via laser ablation with a commercially available, quasi-continuous C O 2 laser system with a central wavelength of 10.6 mm (Epilog Legend 36EXT). Laser ablation increased the roughness (i.e., physical inhomogeneity), resulting in loss of slipperiness in the domains.

Atomic force microscopy (AFM) Surface roughness on SLIC surfaces was characterized using AFM (Bruker Multi-Mode 8-HR) with silicon nitride probes in the ScanAsyst mode. At least 30 images were acquired by scanning 4 3 4 mm areas at a scan rate of 1 Hz and analyzed with Nanoscope Analysis software 1.8 to obtain the root mean square roughness Rrms.

The surface chemistry on SLIC surfaces was characterized using XPS (Physical Electronics PHI-5800 spectrometer). XPS was conducted using a monochromatic Al Ka X-ray source operated at 15 kV, and photoelectrons were collected at a takeoff angle of 45 relative to the sample surface. XPS data were acquired from at least five spatially different locations on the surface, and the spectral analysis was conducted using PHI Multipak software.

Ellipsometry was conducted to measure the thickness of PEG layer on SLIC surfaces at different t sil using a variable angle spectroscopic ellipsometer (VASE-VB-250). A spectral scan of the surface was collected between 500 and 900 nm for an incident angle between 55 and 75 with an increment of 5. The thickness of P EG layer (refractive index = 1.45) was determined using a three-layer planar model (air/ PEG/silica) of the solid surface from the collected spectra. At least five measurements were conducted at different locations on each surface.

Contact angle, sliding angle, and surface tension measurements Contact angles and sliding angles were measured with 20 mL sessile droplets, and surface tension was measured with pendant droplets, using a contact angle goniometer/tensiometer (Rame-Hart 260). A custom-built humidity chamber was used to control the humidity. At least six measurements were performed on each surface at spatially distinct locations. The standard deviation is reported as error with all measurements.

Water condensation experiments were conducted in a custom-built insulated enclosed chamber equipped with an inlet for steam and a transparent glass window. Steam (at 100C and 1 atm) was generated by boiling water in a conical flask and was directed to the inlet of enclosed chamber through insulated pipes for conden-sation.

The substrate was vertically mounted on a Peltier plate set to ambient tem-perature ( z 20C), opposite to the glass window, and it was exposed to steam inside the enclosed chamber. The condensation process was monitored and recorded through the glass window using a camera.

Total internal reflection fluorescence microscopy TIRF microscopy was conducted to study protein adsorption using time-lapse imaging with fibrinogen from human plasma conjugated to Alexa Fluor 647 (Fisher). The time-lapse imaging was conducted using a custom-built microscope 48 equipped with Olympus IX71 body, 100x objective, and a CRISP ASI autofocus system. The excitation was conducted using a 638-nm laser (DL638-328 050, CrystaLaser, Reno, NV). Emission was collected using the appropriate Semrock bandpass filters, and the images were acquired in a water-cooled, back-illuminated E M C C D camera (iXon DU-888, Andor, Belfast, UK) liquid-cooled to 70C with an electronic gain of 60. An enzymatic oxygen scavenging system was used in the imaging buffer to reduce photobleaching. 49 Imaging buffer consisted of 50 mM Tris-HCl (pH 8.0), 10 mM NaCl, 0.15 mg mL 1 glucose oxidase, 34 mg mL 1 catalase, 0.8% (w/v) glucose, and 1% (v/v) b-mercaptoethanol. On SLIC surfaces, time-lapse imaging was conducted at a frame rate of 30 frames hr 1 . On other surfaces, the time-lapse imaging was conducted at a frame rate of 10 frames s 1 . During imaging, the con-centration of fibrinogen was maintained at 5 nM for all experiments.

For our SLIC surfaces, we determined the surface energy "1 = ᵯ 1 + "1 = 22 mN m -1 + 44 mN m -1 = 66 mN m -1 using the Owens-Wendt analysis 5 , with water (polar liquid) and methylene iodide (non-polar liquid) as probe liquids. To separate a compound droplet of water and oil using our SLIC surfaces, water must displace oil from the SLIC surface so that it can slide away while leaving the oil adhered. To determine whether or not water displaces oil from the SLIC surface, we compared the energy of two configurations (see Figures S4A andS4B) -oil displacing water