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Engineering Cell Instructive Materials to Control Cell Fate and Functions Through Material Cues and Surface Patterning Maurizio Ventre, and Paolo Antonio Netti ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.5b08658 • Publication Date (Web): 22 Dec 2015 Downloaded from http://pubs.acs.org on December 30, 2015

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ACS Applied Materials & Interfaces

Engineering Cell Instructive Materials to Control Cell Fate and Functions Through Material Cues and Surface Patterning

Maurizio Ventre, Paolo A. Netti* Department of Chemical, Materials and Industrial Production Engineering and Interdisciplinary Research Centre on Biomaterials, University of Naples Federico II, P.le Tecchio 80, 80125 Napoli Center for Advanced Biomaterials for Health Care@CRIB, Istituto Italiano di Tecnologia, L.go Barsanti e Matteucci 53, 80125 Napoli

KEYWORDS Adhesion, Cytoskeleton; Topography; Material Stiffness; Patterning; Differentiation.

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ABSTRACT

Mastering the interaction between cells and extracellular environment is a fundamental prerequisite in order to engineer functional biomaterial interfaces able to instruct cells with specific commands. Such advanced biomaterials might find relevant application in prosthesis design, tissue engineering, diagnostics and stem cell biology. Because of the highly complex, dynamic and multifaceted context, a thorough understanding of the cell-material crosstalk has not been achieved yet; however, a variety of material features including biological cues, topography and mechanical properties have been proved to impact the strength and the nature of the cell-material interaction, eventually affecting cell fate and functions. Although the nature of these three signals may appear very different, they are equated by their participation in the same material-cytoskeleton crosstalk pathway as they regulate cell adhesion events. In this work we present recent and relevant findings on the material-induced cell responses, with a particular emphasis on how the presentation of biochemical/biophysical signals modulates cell behavior. Finally, we summarize and discuss the literature data to draw out unifying elements concerning cell recognition of and reaction to signals displayed by material surfaces.


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INTRODUCTION Morphogenesis and homeostasis of tissues are the results of intricate and delicate interplays occurring between cells and the surrounding chemical/physical microenvironment. Even small perturbations of such interactions may cause malformations or dysfunctions.1 For a long time the extracellular matrix (ECM) has been regarded as a passive supporting frame in which cells were considered to be the main actors. Today, the ECM is undoubtedly recognized as an active structure, source of the signals that promote, guide and sustain cellular functions. Signals displayed by the ECM can be in the form of biochemical signals (fixed proteins or diffusible factors) mechanical stimuli, (hard/elastic, soft/compliant or gel-like tissues), topographic signals (fibrils, fibers, pores, meshes, protrusions). The effects of soluble factors on cell behavior have been extensively investigated in the past decades and the related studies constitute the foundation of modern experimental cellular and molecular biology. Conversely, the effects of the fixed ECM signals -for instance adhesion sites, topography and mechanical properties- on cell functions are much less known. This raises relevant issues, especially in an in vitro context. In fact, the inherent chemical/physical characteristics of the material substrate for cell cultures cannot be disregarded a priori. For example, cell conditioning with biochemical stimuli might be affected by material properties, which can be relevant when drawing out definitive conclusions on the treatment.2 In particular, properties such as material stiffness, roughness, ligand density and availability, surface charge, hydrophobicity, invariably come in contact with cells and affect their response. In fact, there is growing evidence that this type of signals can be as effective as soluble signals in regulating cell fate and functions.3 Within this context, cell adhesion acquires a central role, since adhesion is a prerequisite for the perception of surface topography, material stiffness and ligand positioning. Additionally, adhesion formation and dynamics dictate the assembly of the actomyosin cytoskeleton, whose contractile forces strongly affect multiple cell functions, such as spreading, migration, proliferation and differentiation.4 The transduction of mechanical signals, i.e. ACS Paragon Plus Environment

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mechanotransduction, either exogenous or generated by the contractile activity of the cytoskeleton, is a new research field that is steadily gaining importance as it holds the promise to provide novel elements to interpret complex biological processes, such as morphogenesis, tissue repair and tumor progression.5 Mechanotransduction relies on the physical interaction occurring at the interface of ECM and cells or, from a biomaterial science perspective, material-cytoskeleton crosstalk. Unraveling such a complex material-cytoskeleton crosstalk would provide novel criteria for designing biomaterial surfaces able to impart specific instructions to cells through adhesive signals at the interface. In this work we review the most relevant findings on material-induced cell responses, with a particular emphasis on cell behavior control through material cues and surface patterning. Throughout the discussion, we highlight the contribution provided by our research group in the field. In particular, we present our recent findings on surface functionalization, cell behavior control with biochemical, mechanical and topographic signals and tissuegenesis control with nanopatterned materials. Finally, we summarize and discuss the massive literature in an attempt to draw out unifying elements concerning cell recognition of and reaction to biochemical and biophysical signals exhibited by material surfaces.

MACROMOLECULES AND DYNAMICS OF CELL ADHESION Cells and materials can interact in various manners according to their relative distance and time of exposure. Material surfaces aimed at interacting with biological entities, like tissues and organs, for a prolonged period, months or years as for prosthesis or scaffolds, are of particular interest in the fields of biomaterials science and tissue engineering. In this context, cells are persistently exposed to the signals displayed by material surfaces, and cell adhesions play a predominant role in interpreting these signals, which affect cell functions. There is growing evidence that cell adhesion need to be finely


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controlled in order to elicit specific cellular responses that ultimately may dictate the performance of devices in situ.6 Transmembrane receptor integrins are widely recognized as the elements that allow cells to adhere to substrates and to perceive a broad spectrum of local chemical and physical properties. In particular, clusters of integrins, known as focal adhesions, enable a firm connection with the substrate through specific ligands, thus mediating the force exchange between the cell and the extracellular space.7 Focal adhesions are not static entities: they form, grow and disassemble via the recruitment, interaction and turnover of several molecular components.8 Integrins belong to a broad class of heterodimeric transmembrane receptors, present in the form of α and β subunits. So far, 18 α and 8 β subunits have been identified. These can interact with each other, thus forming 24 distinct dimers showing various degrees of affinity with ECM motifs. Several amino acid sequences have been found to be specifically recognized by integrins heterodimers. Among these the RGD, YIGSR, YKVAV, LGTIPG, PDGSR, LRE, LRGDN and IKLLI were identified in laminin, RGD and DGEA in collagen I and RGD, KQAGDV, REDV and PHSRN in fibronectin. Generally, the first adhesions to form are located in proximity of the peripheral areas of the cell membrane. Nascent adhesions involve the clustering of very few integrin dimers. They can further grow into larger complexes or disassemble. The fate of the nascent adhesions depends on ligand availability, integrin gathering and stability of the binding between the two. Adhesion growth is not just a matter of cluster size, but it involves the stabilization of the cytoplasmic side with the incorporation of adhesion proteins, such as talin, vinculin, paxillin and focal adhesion kinase (FAK). These proteins also mediate the interaction with and connection to the actin cytoskeleton. In particular, nascent adhesion can engage the retrograde flow of actin through talin. Once coupled, myosin-generated forces can induce a conformational change of talin that exposes additional binding sites for vinculin, which in turn stabilizes and recruits additional proteins.9 Cellgenerated forces are therefore necessary for the maturation of adhesions. Under the action of contractile ACS Paragon Plus Environment

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forces, the maturation process can continue generating focal adhesions (FAs) approximately 1 µm wide and up to 10 µm long. As it has been recently pointed out by Kanchanawong et al., cytoplasmic FA proteins are not randomly distributed, but rather assembled in a multi-ply fashion whose strata fulfill specific functions.10 In particular, upper tails of integrins, FAK and paxillin constitute a signaling layer; talin and vinculin define an intermediate layer with a force transduction role; a top layer of vasodilatator stimulated proteins (VASP), zyxin and α-actinin, that are responsible for regulating the interactions and the assembly of the growing cytoskeleton.

Figure 1. Schematic of adhesion formation and maturation. (A) Nascent adhesion forms after integrin activation and binding with talin. Additional components, such as FAK and paxillin (not-shown) are recruited to stabilize the cluster. New actin filaments engage talin. Myosin contraction stretches talin,


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which exposes binding sites for vinculin. Stable clusters can grow further by addition of other integrins and cytoplasmic molecules, thus generating a focal complex and then a focal adhesion. (B) The multilayered structure of a mature focal adhesion constituted by a signaling layer (integrins, FAK and paxillin), force transduction layer (vinculin and talin) and a linkage to the cytoskeleton (Vasp, zyxin and α-actinin). Owing to the dynamic nature of focal adhesions, it is difficult to dissect the molecular basis regulating adhesion formation and growth with conventional cell culture substrates. Advancements in materials engineering made available surface functionalization strategies that enable a precise control on adhesive signal presentation, in terms of ligand type exposure, density and spatial positioning, i.e. patterning. Different approaches can be pursued to functionalize biomaterial surfaces to show adhesive signals to cells. Small peptides or entire biomacromolecules can be anchored to surfaces through physical (adsorption or precipitation) or chemical (covalent binding) methods.11 Several studies have addressed the issue of precisely controlling adhesion site density to promote cell attachment to material surfaces. More specifically, the role of the surface density of the RGD tripeptide has been widely investigated.12 For instance, the spreading of fibroblasts on glass surfaces requires a minimum RGD density of 1.0×10-15 mol/cm2, which corresponds to a spacing of about 440 nm between peptide ligands, whereas a density of 1.0×10-14 mol/cm2 is needed to promote the formation of focal contacts and stress fibers formation.13 However, the nominal density of the target ligand might not match the actual density, which is effectively displayed to and available at the cell membrane. The differences arise from several factors, such as deposited or incorporated amount, spatial arrangement and accessibility of the peptide, material type and its surface texture. For example, concerning polyethylene terephtalate conjugated RGD, an enhancement of cell adhesion is observed for ligand densities as high as 10-13 mol/cm2.14


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Owing to the importance of ligand engagement and cell adhesion in biomaterials design, the understanding of the influence of the functionalization process and material surface characteristics on ligand presentation acquires a central importance. To this aim, we performed a systematic study to characterize the actual spatial presentation of the RGD signal on polycaprolactone (PCL) and its effect on cell adhesion.15 In detail, we grafted RGD sequences on PCL foils through a two-step procedure: polymer aminolysis to graft functional primary amines on the material surface, followed by a chemical conjugation of the RGD motif. We found that the RGD signal was predominantly located on the boundaries of polymer crystallites at the surface. Furthermore, although the ligand was effectively engaged by cells, RGD penetrated few tens of microns beneath the polymer surface, thus reducing the actual signal availability.15 These data suggest that functionalization strategies must be carefully optimized for each polymer/ligand pair, in order to fabricate biomaterial surfaces that effectively control and guide the adhesion and spreading processes. The vast majority of the works on the presentation of biochemical signals is based on experimental campaigns performed in ‘ideal’ conditions, i.e. well defined chemical environments in which the density of signals is known a priori. This might in principle limit the practical translation of the results in an in vivo context, in which many other bioactive signals may unpredictably overlap with the immobilized ones, thus making it difficult to interpret the outcomes. Consequently, little information is available on the influence of adsorbed protein layers on the signals deliberately engrafted on the biomaterial surface. Along these lines, we investigated the availability of bound vs. adsorbed signals represented by serum-supplemented media and evaluated the effectiveness of ligand covalent conjugation on cell recognition in terms of FA establishment and growth, cytoskeleton organization and cell mechanical properties.16 To this aim, NIH3T3 fibroblasts were cultivated on PCL sheets to which RGD sequences were covalently conjugated through aminolysis-epoxy crosslink reaction. In presence of serum, proteins rapidly adsorbed onto the surface thus masking the bound signal. Even in ACS Paragon Plus Environment

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these unfavorable circumstances, the adhesion machinery of cells eventually selected and attached predominantly to covalently bound ligands (Figure 2A-C). Time course investigations of FA dynamics indicated that cells experienced two different adhesive signals competing at two different time scales: early adhesion, within 4 h post seeding, in which physisorbed proteins readily offer adhesive sites to the cell membrane; late adhesion in which cells ‘dig down’ the physisorbed layer and anchor to bound ligands (Figure 2D).

Figure 2. NIH3T3 fibroblasts cultivated for 24 h in the presence of serum supplemented medium on bare PCL (A); PCL functionalized with RGE (B); PCL functionalized with RGD (C). Note the irregular cell morphology with abundant cytoplasmic vinculin and few FAs and stress fibers when the adhesive signal is only physisorbed. Conversely, thick actin bundles and well defined FAs are visible when the adhesive ligand is covalently linked on the substrate. (D) Putative mechanism of cell digging to establish adhesions on weakly (left) or firmly bound (right) ligand and consequent cytoskeleton buildup. Actin is stained with TRITC phalloidin (red); FAs are stained for vinculin (green). Bar = 10 µm.


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The detailed experimental and image acquisition procedures are described in the Supporting Information. The use of small peptide sequences is not the only mean to functionalize surfaces to control and guide cell adhesion processes. Other biomolecules proved to affect various aspects of cell adhesion from integrin clustering up to cell spreading and migration. Proteoglycans, for example, are known to bind to diverse ECM molecules, which mediate cell adhesion, or to directly interact with membrane receptors that influence adhesion formation. Decorin and biglycan anchored to collagen-based coatings on material surfaces effectively improved FA formation in osteoblastic cells, thus affecting cell adhesion.17 Another class of biomolecules strongly affecting adhesion formation is constituted by growth factors (GFs). These have been widely employed in a soluble form. However, GFs have been shown to affect integrin activation through their cognate receptors and vice versa, GF receptors became active only in presence of integrin clusters bound to ligands.18 Although the intricate molecular mechanisms occurring on the GFs – GF receptors – integrin axis is well beyond the scopes of this article, we wish to emphasize that a dynamic crosstalk between adhesions and GF receptors exists that can result in synergistic effect on common signaling pathways. Therefore a careful selection of the ligands for functionalizing surfaces can enable a fine tuning of cell function by enhancing or suppressing specific intracellular pathways. This notwithstanding, there are comparatively fewer studies exploiting immobilized GFs on surfaces with respect to those dealing with short adhesive sequences. Usually, major challenges in the chemical conjugation of GFs to synthetic surfaces is represented by the preservation of GF bioactivity and the correct presentation to receptors. However, advancements in chemical functionalization strategies enabled the fabrication of material platforms displaying complex patterns of GFs and ligands.19 Shahal et al. functionalized glass with biotinylated RGD and/or Epidermal Growth Factor (EGF) coupled with biotinylated PEG brushes via the biotinNeutrAvidin interaction.20 They investigated the adhesion process of A431 cells when exposed to ACS Paragon Plus Environment

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various combination of ligand and GF densities. In general, bound EGF was more efficient in affecting cell adhesion with respect to the soluble form. Interestingly, EGF improved adhesion with the formation of FA and mature stress fibres up to a certain threshold level, above which EGF promoted detatchment. The authors interpreted this as a mechanical unbalance at high EGF density, which causes FA to not withstand the high level of contractility induced by the GF. Altogether these data confirm the efficacy of chemical bioactivation of material surfaces. This might be relevant to study adhesion processes or cytoskeleton assemblies in complex environments or to exploit mechanotransduction events in contexts that more closely mimic an in vivo situation.

PATTERNING SIGNALS ON MATERIAL SURFACE The processes of FA formation and maturation, along with cytoskeleton assembly require the spatiotemporal orchestration of molecular events (ligand engagement, integrin activation, cytoplasmic proteins recruitment, actin network build-up).8 Functionalization of surfaces with uniform densities of ligand is not sufficient to modulate the complex molecular interplays regulating cell adhesion. A better understanding of the dynamics of FAs and cytoskeleton requires harnessing more refined techniques that enable a precise control of ligand positioning. In what follows, we briefly review some of the technologies that became popular in the biotechnological sector for the development of materials aimed at gaining a tight spatial control on the adhesion events. Advancements in microelectronics and optics promoted the development of processing techniques that allowed the fabrication of micro- and nano-structured materials in a highly consistent manner. Among the various techniques, photolithography is undoubtedly at the foundation of micro- and nanofabrication technologies, currently used to create patterned surfaces, mainly molds and masters, for ACS Paragon Plus Environment

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biotechnological and biomedical applications. In the process, UV light transfers a pattern from a photomask to a light-sensitive photoresist that is subsequently developed to generate the desired substrate. Soft lithography is a complementary extension of photolithography. This technique relies on the coating of a liquid precursor, usually polydimethylsiloxane (PDMS), on a photolithographygenerated master. The precursor is chemically or thermally cured, thus producing an elastomeric patterned stamp. Soft lithography does not refer to a single technique, but it is rather a collection of techniques in which the soft stamp is used for printing, embossing and molding. Among these techniques, microcontact printing (µCP), replica molding (REM), micromolding in capillaries (MIMIC) are amongst the most popular in the biotechnology field. In µCP, the elastomeric stamp is used to transfer bioactive molecules on a surface, usually a self-assembled monolayer (SAM) of polyethylene glycol (PEG). In this case the non-inked PEG displays no bioactivity, which is confined to the inked regions. In REM, the patterned elastomeric stamp can be used as it is or as a master to fabricate patterned substrates. MIMIC is based on the spontaneous filling by capillary suction of a fluid containing molecules of interest between two surfaces in conformal contact. The solutes in the fluid can then be adsorbed on one surface or can be chemically or physically treated to crosslink, thus creating a microstructure replicating the shape of the channel network. Despite the outstanding simplicity in using soft lithographic technologies, the lateral resolution of patterns is usually limited by the diffraction of the UV light used to fabricate the master. Resolution of ~ 1 µm can be easily achieved, whereas narrower features require special apparati; for instance, a feature size of ~ 300 nm can be achieved with extreme UV light. To overcome such limitations, more refined techniques aimed at challenging the diffraction limit of light were developed. Among these, electron beam lithography (EBL) and focused ion beam lithography (FIB) allow the fabrication of features with size in the order of tens of nm. The working principle is analogous to the conventional photolithographic technique. However, these techniques do not require a mask, as the beam path is ACS Paragon Plus Environment

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controlled with electromagnetic lenses. Furthermore, FIB allows atoms to be displaced from or deposited onto the material surface, in which case it is possible to achieve subtractive or additive lithography on the final substrate directly, without further development. If the pattern is created on a hard substrate, this can be employed to mold a viscous polymer (i.e. a polymer above the glass transition temperature or a photocurable one), which is subsequently stabilized via cooling or curing, in a process known as Nanoimprint Lithography. Other technologies to nanopattern materials for biotechnological applications, such as dip-pen lithography, involve the use of the atomic force microscope (AFM) equipped with specifically designed tips and cantilevers coated with biomolecular solutions (inks).21 Such tips act as pens that write nanopatterns on surfaces by depositing the ink in a controlled fashion. Careful optimization of the processing conditions enables the creation of 10 nm features. Although these techniques allow a large increase in the spatial resolution of pattern features, they suffer from intrinsic high costs and are timeconsuming; therefore, they are not the preferred techniques for patterning large areas. To overcome the above-mentioned problems, several thermodynamic-based techniques have been developed. Among these, block copolymer self-assembly and polymer demixing were extensively used to create patterned substrates to control cell behavior in a simple, rapid and cost-effective manner. The techniques are based on the aggregation or segregation of polymers. For example, in polymer demixing, phases are segregated from spin-coated blends upon solvent evaporation. This processes lead to the formation of a topographic pattern with features having a varying lateral dimension but consistent depth. Controlling the composition and thermodynamic properties of the system it is possible to obtain some degree of control on the topographic feature. However, patterns thus produced rarely exhibit long-range order.


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To gain a better control on the feature size, Glass et al. developed a technique, named micellar lithography, which easily patterns large surfaces without using expensive equipment.22 More specifically, the technique is based on the even arrangement of polymeric micelles, containing metallic particles in their core, on flat surfaces. The hybrid organic/inorganic micelles form a closely packed coating in which interparticle distance is dictated by micellar dimension. With this technique it is possible to fabricate quasi-hexagonal patterns of metallic particles with spacing in the 10-200 nm range, depending on the processing conditions. In the field of inorganic materials, ordered assemblies of titania nanotubes were successfully produced via an anodization bottom-up.23 Such a technique allows to create vertical, densely packed tubes, with a tight control on tube diameter: from tens up to hundreds of nanometers. Practical applications of some of the technologies here reported, for the fabrication of platforms to affect cell functions through adhesion control, will be discussed in the following chapter.

USING MICRO- AND NANO-PATTERNING TO DISSECT CELL-MATERIAL INTERACTIONS Micro- and nano-patterning technologies combined with surface functionalization strategies, allow to gain a great control over a broad range of cell behaviors. For instance, nanopatterning may control adhesion formation at a single integrin dimer level, whereas micropatterning can confine cells thus dictating their shape. This approach is fundamental to study complex cell behaviors in a chemicalphysical defined environment, which can ultimately help in dissecting the role of adhesion, cytoskeleton and cell shape on cell functions. In what follows, we briefly review the most relevant findings on material-induced cell behaviors such as adhesion, migration and differentiation.


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Adhesion control through nanoscale functionalization. In principle, cell adhesion can be controlled by modulating positioning, clustering and density of specific ligands, which can be proteins or peptide sequences. Of course, proper spatial control of adhesion formation and growth requires that adhesion promoting zones are separated by the repelling ones. The development of specifically functionalized SAMs, polymer brushes and block copolymers have largely contributed to achieve both a sharp segregation of adhesive/non-adhesive regions and a nanoscale control on ligand positioning.24 Micellar lithography also proved to be particularly suitable for cell adhesion control, as it is sufficiently versatile to enable modulation of feature size, i.e. lateral spacing and cluster size. Such technological approach allowed answering relevant questions related to adhesion formation and maturation. Arnold et al. demonstrated that adhesion formation requires an interligand spacing, i.e. an RGD conjugated to a gold nanoparticle, of ~ 70 nm at most.25 Patterns with small interligand spacing (i.e., high ligand density), caused cells to adhere firmly on the surface, displaying well-defined FAs and contractile stress fibers. Conversely, cells seeded on patterns with high interligand spacing (i.e., low ligand density) were less spread, round, with small FAs. By exploiting self-assembly of block Polystyrene- Poly(ethylene oxide) copolymers, George et al. were able to control the nanoscale presentation of adhesive ligands (RGD) in a precise and rapid manner.26 Their elegant method allowed the fabrication of random patterns in which ligand separation ranged from 44 to 62 nm. According to the data of Arnold et al.25 they found improved fibroblast spreading on substrates with high RGD density. Owing to the dynamic nature of FA formation and growth, ligand spacing is not the only parameter affecting adhesion formation and growth. Integrin clustering is important to determine whether a FA is sufficiently stable to permit cell spreading. With the aid of EBL, Schvartzman et al. fabricated arrays of metallic nanoclusters whose configuration varied from dimers to heptamers with each element containing a single RGD.27 The authors demonstrated that a tetrameric clustering of ligands was necessary to enable cell spreading. These findings are in close agreement with a previous study by ACS Paragon Plus Environment

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Maheshwari et al. who exploited star-PEG hydrogels functionalized with YGRGD ligands.28 By carefully controlling the functionalization process, one to nine RGD ligands were selectively attached to the star-PEG arms with an intercluster spacing varying from 6 to 300 nm. Generally, clustered ligands were more effective in allowing cell spreading compared to individual ligand display. For instance 3.0 × 105 ligand/µm2 were less effective in inducing cell spreading than 2.3 × 103 ligand/µm2 with ligands assembled in clusters of nine. Collectively, these data demonstrate the extraordinary sensitivity of the cell to nanoscale positioning of adhesion sites: small variation in ligand spacing or clustering may result in either firm adhesion or detachment. Effects of material features on adhesion proliferation and migration. The segregation of cell adhesive/repellent zones at the nanoscale is essential to alter adhesion events at the molecular level. However, the continuous variation in ligand densities, for example in the form of one- or twodimensional gradients, profoundly affects different aspects of cell behavior, such as spreading, polarization and migration. Several chemical and physical methods were developed to generate concentration gradients of ligands on synthetic substrates, with a precise control on gradient slope and average concentration.29 Platform thus produced can be useful to investigate the behavior of cells experiencing continuous local changes in ligand density. Generally cells polarize and migrate towards the higher concentration of the ligands. Simth et al. produced a gradient of fibronectin starting from graded SAMs obtained via diffusion of alkanethiols on gold. Endothelial cells plated on the graded substrate showed high migration speed with steeper gradients.30 Results in agreement with these were obtained by our group, by generating RGD gradients on a PEG hydrogel surface by means of a fluidic device.31 Mouse fibroblasts showed high speed values when seeded on graded hydrogels than that measured on gels with uniform distribution.


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Cell adhesion, migration and proliferation have also been widely investigated on micro- and nanostructured material surfaces. While the effects of large micropatterns, i.e. feature size comparable with cell size, may result in a physical confinement of the cell body, topographic patterns with feature size similar to that of the sensorial organelles, such as integrins, FAs, filopodia and lamellipodia, exert a potent effect on the processes of surface recognition and adhesion. For instance, cells are generally observed to elongate and migrate along topographies constituted by straight and parallel channels and such an effect is enhanced on narrower ridges (< 10 µm).32 The recognition and response to topographic features involve different processes of the cellular adhesion and contractile machinery. In a recent publication we highlighted the importance of filopodial probing and FA establishment and growth in the response of preosteoblasts to parallel microgratings in PDMS.33 These were designed in order to facilitate filipodial probing or adhesion growth. Specifically, 5 µm wide ridges with a lateral spacing of 5 µm, did not allow filopodia to establish additional FAs on the ridges next to the cell sides since average length of filopodia rarely surpasses 5 µm. Therefore, cells predominantly migrated on a few ridges in a back-and-forth fashion. Conversely, on patterns of 2 µm wide ridges with a lateral spacing of 2 µm, cells were able to form transverse processes that allowed them to move in directions different from that of the pattern, in a sort of random migration. In this case, however, unidirectional migration was restored by performing oxygen plasma treatment, which improved FA maturation and growth. The length of the FAs orthogonal to the pattern direction was limited by the ridge width, as ‘suspended’ FAs were not observed. Such a limitation caused the orthogonal FAs to collapse because they were unable to withstand the high traction forces generated by the actomyosin cytoskeleton. Conversely, the majority of mature FAs were those growing along the pattern direction and this caused the cells to align and migrate along the pattern direction. Our data confirmed that topographies can be seen as alternating zones conducive or unfavorable to FA formation


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and growth and that additional surface treatment can amplify or depress such an adhesion contrast. Therefore, topographic patterns are assemblies of surfaces that are readily accessible for lamellipodia and filopodia, top of ridges or pillars, as well as impervious recesses like the bottom of grooves and pits. The accessibility of a given region to cell membrane protrusions and receptors is dictated by the geometric characteristic of the pattern, i.e., feature depth and pitch. When these parameters are such that cell membranes cannot accommodate surface recesses, the cell is suspended on the top of the pattern and adhesion is limited to specific parts. In this particular situation, the limiting value of feature depth below which the topographic signal ceases its regulatory effect on adhesion formation was reported to be 35 nm, approximately.34 The combined effect of ridge-groove-depth size on osteoblast orientation and migration was carefully analyzed by Lamers et al.35 The authors reported that ridges with lateral size < 75 nm allowed integrin bridging the grooves, therefore clustering, firm adhesion and low motility. A change in cell response was observed when lateral spacing exceeded 75 nm. In fact, cells were highly aligned in patterns with ridges and grooves of 80 nm width or on patterns with ridges of 50 nm and grooves of 100 nm. Conversely, no significant orientation was observed on the inverse pattern, i.e. ridges of 100 nm and grooves of 50 nm. Furthermore, feature depth dramatically affected cell orientation as patterns with depth < 35 nm were much less effective in coordinating cell alignment. An elegant study by Kim et al. provided deeper insight into the cell’s sensitivity to feature spacing.36 To this aim, a substrate with graded topography (ridge width 1 µm and height 400 nm and lateral spacing continuously varying from 1 to 9 µm) was developed through EBL. NIH3T3 fibroblasts cultivated on this substrate displayed enhanced alignment on the narrower topography, whereas cell speed was optimal at the intermediate ridge spacing. Additionally, trajectories seemed to point towards a common location in the intermediate ridge spacing. Altogether, these data demonstrate an exquisite sensitivity to micron-scale variations of pattern features and possibly that cells tend to migrate towards an optimal zone.


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The guiding effects of topographic patterns on adhesion and migration arises from alignment and spatial distribution of FAs caused by the confining effect of the topographic features. Ordered arrays of FAs cause the cytoskeleton to assemble in the form of parallel actin bundles that force the cell to be elongated and oriented along the pattern. A better understanding on the influence of topographic patterns on the dynamics of FA formation and cytoskeleton assembly could provide elements to conceive patterns that induce the cytoskeleton and the cell body to acquire shapes defined ab initio, thus altering cell behavior more effectively. Along this line, we have recently investigated the dynamic interplays between FAs and actin bundles on nanopatterned PDMS surfaces, constituted by parallel and straight channel 700 nm wide.37 Time course experiments of cells expressing fluorescent paxillin and actin revealed complex structural remodeling events during cell recognition and reaction to nanopatterned surfaces. First, we observed very peculiar cytoskeleton assemblies, which spontaneously formed on the nanopatterns. Basal actin fibers, orthogonal to the pattern direction (Figure 3A), formed in the early adhesion phases and remained so even at longer culture times. These were scarcely sensitive to myosin II contraction, suggesting that such a structure might have a stabilizing role under the nucleus rather than a contractile function. Furthermore, time course experiments revealed that this configuration of the actin network was kinetically favored with respect to the arrays of thick fibers located in the apical part of the cell (Figure 3B). Basal fibers formed a sort of stabilizing scaffold underneath the nucleus. Second, actin fibers not directed along the pattern direction usually terminated with dashed adhesions located in proximity of adjacent ridges (Figure 3C). These proved to be unstable, as the dashed adhesions collapsed under the effect to contractile forces. Thus, only FAs along the pattern were able to withstand cell contractility. Such an FA remodeling had an obvious effect on cell shape, which was elongated along the pattern, but also a less intuitive effect on nuclear shape. In fact, orderly arrays of actin fibers exerted a lateral compression force on the nuclear envelope in a more efficient manner with respect to what observed with disordered fibers (Figure 3D). This was mainly


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caused by the confining effect exerted by the nanopattern on the FAs as these were not able to slide orthogonally to the pattern ridges to accommodate actin fiber straightening. Nanopatterned adhesive signals can transfer mechanical information up to the nucleus effectively and rapidly, which might result in a modification of the transcription events (see below). Therefore, nanopatterns can be used to modulate the stress state acting on the nucleus.

Figure 3. Focal adhesion and cytoskeleton assemblies of MC3T3 cells, cultivated on nanopatterned PDMS substrates for 24 h. Basal (A) and apical (B) configuration of actin stress fibers. Dashed FAs on adjacent ridges connected by individual stress fibers (C). Contraction of actin fibers wrapping the nuclear envelope around causes extensive squeezing of the nucleus (D). Actin is stained with TRITC-


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phalloidin (red); FAs are stained for vinculin (green); Nuclei are stained with DAPI (blue). Pattern direction is parallel to the horizontal axis. Bars = 10 µm. The detailed experimental and image acquisition procedures are described in the Supporting Information. Concerning cell proliferation, generally nanotopographic patterns have been reported to reduce proliferation, if compared to flat surfaces.38,39 However, there is also evidence of increased proliferation for certain cell types, whereas no obvious trends have been reported so far on the effects of patterns in the form of pits and pillars.40 Altogether these data suggest that the mechanism regulating topographyinduced cell proliferation are much more complex than those for adhesion and migration and might be dependent on cell type, pattern features and culturing conditions. The effects of substrate stiffness on cell behavior have been widely documented. From the pioneering work of Harris et al.41 who first documented how cell-generated forces could deform the underlying material substrate, a multitude of investigations has been performed and has demonstrated that cell contractility is essential in probing the stiffness of the extracellular microenvironment. This, in turn, modulates responses such as adhesion, proliferation and migration. Soft materials (0.1 – 1 kPa in modulus as soft hydrogels) depress focal adhesion maturation and adhesions remain small and characterized by a high turnover. Conversely, cells cultivated on stiff materials (10 – 100 kPa) display extended FAs with an increased level of tyrosine phosphorylation of FAK and paxillin expression.42,43 These observations, along with others, confirmed that stiffness sensing elicits specific intracellular signaling events. Similarly, cytoskeleton assembly is dramatically affected by substrate stiffness. Yeung et al. observed that fibroblasts cultivated on soft polyacrylamide (PAM) gels (~ 0.2 kPa) were nearly spherical with no detectable stress fibers, whereas on stiffer gels (~ 3.6 kPa) cells were oblong with a clearly visible cytoskeleton.44 A similar response was observed in other cell types. As expected, stiffness-mediated assembly of FAs and cytoskeleton also induced modification in the migratory


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behavior of cells. The relationship between material stiffness and cell speed is not linear and depends also on other factors, such as ligand density; generally, however, higher migration rates are observed on compliant gels, whereas cells tend to be stationary on stiff materials.45 Interestingly, many cell types, if cultivated on substrates displaying gradients of stiffness, have the tendency to migrate towards the regions in which the material is more rigid, a phenomenon conventionally referred to as durotaxis.45 By using a fill-molding in capillaries (FIMIC) technique Diez et al. fabricated PEG based hydrogels with patterns constituted by straight channels of mismatching stiffness.46 More specifically, a microgrooved stiff PEG hydrogel was prepared by REM and UV curing. Then, with the mimic technique, the grooves were filled with a second formulation of PEG precursors. After curing, the softer second hydrogel crosslinked and attached to the stiff master. Cell adhesion and migration were predominantly observed on the stiffer regions and investigations with the scanning electron microscope revealed a negligible surface topography. Therefore, the mechanical pattern acted in an analogous manner to a biochemical pattern, in which adhesive spots are engrafted in a repellent background. Most of the works dealing with the influence of the mechanical properties of the substrate on cell adhesion are based on hydrogels or elastomers, chiefly PAM, PEG, PDMS. Usually, a single material parameter, namely the stiffness, is used to correlate cell response with mechanical properties. While this can be reasonable for linearly elastic materials, it can be inadequate to define the mechanical behavior of non-linear materials or viscoelastic materials. In fact, certain formulations of soft hydrogels and elastomers used in mechanobiology studies, are viscoelastic in nature. Thus, the hypothesis of linear elasticity can be misleading when drawing out general conclusions on the role of material mechanical properties on cell behavior, especially when time-dependent effects are taken into consideration. This issue was addressed by Cameron et al. who cultivated human Mesenchymal Stem Cells (MSCs) on different PAM hydrogels, displaying the same elastic modulus but different loss moduli.47 They observed increased cell spreading and proliferation, but a decreased size and maturity ACS Paragon Plus Environment

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of FAs, for cells on hydrogels with the highest loss modulus. Apparently, the decrease in isometric tension (i.e. contractility) is compensated by an increase of isotonic tension, which favors spreading and more dynamic FAs. Thus, cells are able to perceive different aspects of the mechanical properties of the environment, other than stiffness alone, that strongly affect cell morphology and functions. It has to be pointed out that, beside the effects caused by substrate viscoelasticity, the ‘chemistry’ of the underlying material can influence cell response to mechanical cues. To bridge this gap, we have recently reported NIH/3T3 fibroblast response in terms of time dependent adhesion, spreading and migration to different PAM and PDMS substrates with stiffness ranging from 3 to 1000 kPa.48 In particular, soft PAM behaves as a virtually elastic material, whereas stiffer PDMS has a non-negligible viscous component. Substrates were engineered in order to provide cells with the same ligand density, wettability and roughness, thus keeping material ‘chemistry’ and stiffness as the only two discriminants. Generally, our results confirmed those of other reports, i.e. longer FAs and slower cell migration on the stiffer substrates. However, we also observed that FAs act as mechanosensors, as their morphometric characteristic were sensitive to material stiffness as soon as 6 h post seeding. Additionally, we found a general increase in cellular mechanical properties, as measured with multiparticle tracking, suggesting that part of the outside-in signaling also involved reorganization of the cytoskeletal components that resulted in a stiffer cytoskeleton. While these properties well correlated with substrate stiffness regardless of the material type, the cell spreading area followed a different trend, i.e. positive correlation with the stiffness of the elastic PAM but no correlation with the stiffness of the viscoelastic PDMS. This suggested that on the adhesion level cells respond to the elastic properties of the substrate only, whereas on a larger scale, i.e. spreading, viscous components cannot be neglected. Stem cell fate regulation through material features. Adhesion and cell-generated forces are two intimately connected phenomena, essential for cell survival as well as for allowing cells to fulfill vital ACS Paragon Plus Environment

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tasks as proliferation and migration. So far, we have presented numerous examples demonstrating the profound sensitivity of these functions to extracellular signals. In vivo cell adhesion and contractility play a crucial role in sculpting tissues and organs during embryogenesis. However, they also proved to regulate signaling pathways, gene expression and cell functions.49 The potency of cell adhesion and contractility in regulating cell fate and functions was clearly demonstrated by Chen et al. in 1997.50 In this landmark study, the authors cultivated endothelial cells on adhesive islands of various dimensions fabricated with µCP. Surprisingly, apoptosis and proliferation correlated with the extension of the adhesive islands or, more precisely, with the cell area. In particular, smaller islands provoked high apoptosis, whereas patterns that induced extensive spreading favored proliferation. This work stimulated the attention toward the comprehension of the effects of cell adhesion, most notably cell shape, on cell fate. Owing to their extraordinary plasticity, stem cells are particularly sensitive to extracellular signals since in vivo they are constantly exposed to a multitude of biochemical and biophysical signals provided by their niche. Decisions on self-renewal/commitment are based on the spatiotemporal integration of these signals. McBeath et al. proved that stem cell fate is highly sensitive to cell shape changes in vitro.51 By cultivating human MSCs on micropatterned adhesive islands, they proved that the cell area was a determinant of cell differentiation as small islands promoted adipogenesis, whereas larger islands osteogenesis. Furthermore, the authors suggested that cell contraction was involved in lineage specification through the RhoA/ROCK pathway, as direct manipulation of RhoA modulated adipogenesis/osteogenesis. Even in presence of osteogenic media, RhoA negative cells became adipocytes and vice versa: consistently active RhoA induced osteogenesis in presence of adipogenic media.


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Following studies refined the comprehension of the molecular mechanisms that regulate cell shapeinduced differentiation. Kilian et al. created adhesive patterns through µCP able to induce growing levels of contractility in human MSC, i.e. blunt curves vs. sharp corners and rectangles with different aspect ratio (AR).52 They suggested that adhesive shapes promoting myosin-driven contractility, such as pointed stars or stretched rectangles, enhanced osteogenesis, whereas adhesive shapes in the form of round flowers or small squares induced little cell contractility and promoted adipogenesis. These findings were confirmed by Peng et al. who found a direct correlation between cell perimeter and osteogenesis, as high values of cell perimeter implied convoluted cell borders and -hence- a higher contractility.53 Conversely, cell perimeter negatively correlated with adipogenesis, confirming previous observations.51 Additionally, they extended the investigation concerning the effects of rectangular ARs with differentiation and found that osteogenesis is not monotonically correlated with AR, suggesting that osteogenesis is obtained at intermediate levels of contractility. Even though these works aimed at highlighting the correlation between cell shape and cell fate, the conclusions were drawn out from observation of cell response in presence of induction (or coinduction) media. A question may arise on whether cell shape is sufficient to guide lineage specification without the aid of exogenous biochemical signals. Along these lines Yao et al. performed similar experiments, i.e. culture of rat MSCs on adhesive rectangles of different AR in absence of induction media, for an extended time interval.54 They found that cell shape still was a strong effector of stem cell differentiation even in the absence of biochemical stimulation. However, peaks in differentiation, adipogenesis/osteogenesis, were found at later time points (19 days) with respect to those observed in presence of induction media (7 days). Since micro- and more effectively nano-topographic patterns, as well as substrate stiffness, affect FA formation and dynamics and therefore cytoskeleton assembly and contractility, it is expected that


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topographic and mechanical cues can regulate stem cell fate and functions. Indeed, Dalby et al. demonstrated that disordered, but not random, arrays of nanopits with lateral spacing of approximately 300 nm improved osteogenesis of human MSC in basal media and with an efficiency similar to that produced by biochemical stimulation.55 The same group later reported that a perfect square lattice of the same nanopits was able to maintain human MSC self-renewal up to 8 weeks.56 Altogether, these data demonstrate that topographic patterns that allow FA assembly, thus stabilizing intracellular tension, promote osteogenesis. Conversely, patterns that induce the formation of FAs sufficiently large to allow proliferation, but below a certain threshold that triggers the activation of differentiation pathways, are requested to maintain multipotency. These studies are accompanied by others that further investigated the role of topographic patterns in stem cell differentiation. Of particular interest in the context of regenerative medicine is the evidence of human MSC transdifferentiation into a neuronal lineage induced by micro- and nano-gratings. In fact, Yim et al. reported increased expression of neuronal markers on patterned gratings with a ridge dimension of 0.35 – 10 µm and ridge to groove ratio of 1:1.39 Such an effect was enhanced with the use of retinoic acid and on narrower gratings. Similarly, Lee et al. demonstrated that submicrometric gratings induced neurogenesis of human Embryonic Stem Cells (ESC) with a high efficiency and without biochemical induction.57 The demonstration that the mechanical properties of the culturing substrate can direct stem cell fate was first provided by Engler et al. in 2007.58 They cultivated human MSCs on PAM hydrogels that had a stiffness comparable to that of brain (0.1 – 1 kPa), muscle (8 – 17 kPa) or bone (25 – 40 kPa) and observed that the different substrates promoted neurogenesis, myogenesis and osteogenesis, respectively. More recent studies have also demonstrated that undifferentiated cells cultivated on specific ranges of stiffness can undergo self-renewal. For instance, Chowdhury et al. cultivated murine


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ESCs on soft PAM hydrogels (0.6 kPa) and even in the absence of Leukemia Inhibitory Factor, murine ESC maintained the ability of forming undifferentiated and homogeneous colonies of up to 15 passages.59 In a different study, Gilbert et al. reported that murine Muscular Stem Cells (MuSCs) selfrenewed when cultivated on PEG hydrogels that had comparable stiffness to that of natural muscle (~ 12 kPa). Renewed MuSCs also displayed superior regenerative properties if transplanted in vivo.60 These data suggest that the maintenance of self-renewal is best achieved when the mechanical properties of the culturing substrate match those of the original niche, whereas culturing of undifferentiated cells on rigid plastic/glass leads to heterogeneous commitments and fluctuation in the expression of pluripotency genes.61 It is self-evident, at this point, that -regardless of the signal shown by the substrate, whether topographic, mechanical or biochemical- surfaces that promote contractility via adhesion and spreading are inductive for an osteoblastic phenotype. Conversely, surfaces that depress, but do not annihilate, adhesion and spreading guide stem cells towards less contractile phenotypes such as adipocytes and neurons or even maintain self-renewal. Therefore, common molecular mechanisms that regulate stem cell fate should exist regardless of the underlying biochemical/biophysical signal. These mechanisms occur at different levels from adhesion molecules up to contractile structures and presumably mutual interactions occur in between. Exogenous signals in the form of topographic or biochemical patterns, as well as substrate stiffness, all affect integrin activation and clustering. As previously noted, many other signaling proteins participate in the formulation of focal complexes/adhesions. Some of these molecules are implicated in numerous pathways eventually dictating cell fate and functions.4,5 Adhesion dynamics profoundly affects the RhoA/ROCK pathway, whose activity regulates the osteogenic/adipogenic fate of MSCs.51


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Cell-generated contractility might also impact at the transcriptional level directly. In fact, nuclear lamins, proteins forming the nucleoskeleton, are connected to the actin cytoskeleton through the LINC complex. Therefore, cytoskeleton stresses can be transferred to the nucleus causing a structural modification of DNA sequences, thus facilitating binding of transcription factors.62 Additionally, accessibility of DNA sequences to transcription factors is regulated by DNA acetylation/deacetylation. Interestingly, topographic signals proved to affect these processes by impacting on nuclear shape and histone acetylation patterns.63 Great care should be paid when attempting to translate these findings in a three dimensional context. Recent evidence suggests that “dimensionality” is a biophysical cue that strongly affects stem cell fate and functions. Huebsch et al. reported that murine MSC encapsulated in a physically crosslinked alginate hydrogels with varying stiffness, underwent fate decisions that were uncorrelated from cell spreading and morphology, in clear contrast to what observed in two dimensional contexts.64 Rather, hydrogel elasticity regulated integrin clustering, ligand reorganization and therefore cytoskeletal tension, which correlated with osteogenesis. Thus, cells perceive mechanical cues as different modes of adhesion-ligand presentation. More recently, Khetan et al. have added a crucial piece of information in this concept by emphasizing the importance of degradation in affecting human MSC lineage specification.65 Using hydrogels with independently tunable stiffness, adhesivity and degradability, the authors remarked that the cell ability to degrade and interact with the 3D matrix is fundamental to generate tension and assemble FAs. This eventually dictates fate decision, regardless of cell spreading and morphology. These data undoubtedly demonstrate the powerful control that biochemical/biophysical signals can exert on cells: from simple adhesion up to gene expression. Practical applications of material-induced stem cell commitment might include the fabrication of culturing devices to expand undifferentiated


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populations of stem cells to be used in cell therapies, or the design of prosthetic surfaces able to improve implant integration and performances in vivo. Nevertheless, we believe that the potential of material features have yet to be fully explored. In fact, the vast majority of the research on the subject generally deals with isolated cells, whereas the effects of material features on dense stem cell populations is largely unknown. Even if not directly related to undifferentiated cells, a few examples demonstrate that micro- and nano-patterning can effectively impact on the organization and functions of tissues grown in vitro. ECM rich tissues with an ordered structure were obtained by Guillemette et al. who demonstrated that elastomeric microgratings with a 4 µm periodicity induced corneal fibroblasts to form a collagen rich multilayer, whose microarchitecture was analogous to the lamellar structure of native cornea.66 Using a PEG based nanograted substrate, Kim et al. produced myocardial tissues in vitro. The authors found an optimal combination of ridge/groove dimensions on which the tissues that developed displayed improved electrophysiological characteristics and a higher expression of gap junction markers with respect to tissues grown on flat PEG.67 These data demonstrate that tissues grown in vitro possess a high level of sensitivity towards nanoscale cues and that topographic patterns might exert a powerful control on tissue architecture and functions, consequently. Human Pluripotent Stem Cells (PSCs) possess an intrinsic ability of self-organizing in vitro, producing differentiated and functional organoids. By modulating the culturing conditions, PSC aggregates produced intestinal, renal, cerebral or optic cup organoids.68 Their cultivation relied on conventional 2D substrates, growth in suspension or encapsulation in Matrigel. Yet, there is evidence that contractility is essential for driving specific morphogenetic processes in vitro.69 Based on these observations, we hypothesized that nanopatterned surfaces could guide the self-organization process of human MSCs in vitro and -possibly- tissuegenesis.70 Human MSCs were cultivated on nanograted PDMS substrates with 700 nm wide ridges, 250 nm high (Figure 4A). In 1 week cultures, cells selforganized in the form of 3D long and thin stripes invariably oriented orthogonally to the pattern ACS Paragon Plus Environment

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direction (Figure 4B). The supplementation of the medium with ascorbic acid promoted the synthesis of collagen and thickening of the structure, consequently (Figure 4C-D). Gross morphology of cell structures was reminiscent of tendon histology. Indeed, microscopic and ultrastructural investigations provided evidence of collagen assembling processes similar to those occurring in embryo tendons.

Figure 4. Nanopattern mediated tenogenesis. MSCs cultivated on nanograted PDMS adhere and elongate along the pattern direction at 24 h post seeding (A). After 7 days of culture, human MSCs form supracellular aggregates directed orthogonally to the pattern direction (B). Supplementation of ascorbic acid, for 7 additional days, enables the formation of thick and long cylindrical tendon-like structures (C). Multiphoton imaging reveals a bright second harmonic generated signals produced by


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densely packed collagen fibrils (D). Pattern direction is parallel to the vertical axis. Actin is stained with TRITC phalloidin (red); cell nuclei are stained with DAPI (blue). Bar=50 µm. The detailed experimental and image acquisition procedures are described in the Supporting Information. Tenogenic induction was confirmed by gene expression analysis that showed upregulation of tendonspecific and associated markers. Time lapse observation of the MSC self-organization and tendon-like structure formation revealed that cells coaligned to the nanopattern generated contractile forces. Such a spatial coordination resulted in a guided self-organization, which defined the shape of the supracellular structures. Interestingly, an intermediate level of adhesion was required to enable tissuegenesis, since enhancement or depression of cell adhesion led to the generation of cell multilayers or irregular spheroids, respectively. Therefore, not only the direction of the nanopattern, but also the initial adhesion conditions need to be precisely tuned to guide MSCs towards the correct process of tenogenesis. Our observations disclose additional powers of the surface nanopatterning, since it can effectively control MSC behavior up to tissuegenesis. This opens up new strategies to design material surfaces able to guide tissuegenesis events in vitro; in principle, by modulating pattern features and adhesivity, it is possible to control cell self-organization, tissue architecture and functions.

FUTURE DIRECTIONS IN SURFACE ENGINEERING FOR BIOMEDICAL APPLICATIONS The examples on material-induced cell responses allow us to predict, with a certain confidence, that signal patterning to control adhesion and contractility will soon find practical employment in clinical contexts or relevant biotechnological applications. However, to achieve this final goal, additional and specific studies need to be carried out. First, we are still unable to predefine a precise set of material cues that elicit specific cell responses. In detail, we have not yet developed a thorough understanding of


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cell-material interactions so that we can specifically relate material features with intracellular pathways and -ultimately- gene expression. Since it is technically difficult, if not impossible, to totally isolate different types of signals out of materials, i.e. each and every material invariably possesses a mechanical property, a surface chemistry and a topography, it is more likely that cellular responses are the result of an integration of multiple signals acting simultaneously. This makes the unraveling of the effects of material features on cell behavior arduous. The development of devices for the highthroughput screening of a broad array of material features acting synergically could in part solve this problem. Such an overwhelming task might allow to correlate sets of material features with cellular responses empirically. However, this might not be sufficient for a deep understanding of the governing mechanisms that occur during cell-material interactions. Therefore, high-throughput analyses must be supported by refined molecular investigations acting in real-time. To this aim, gene reporting techniques and super resolution live microscopy could be particularly useful. Second, cells and more importantly stem cells are subjected to time-changing signals in vitro. Different temporal presentation of the same signal might elicit very different responses. This is a well-known issue when dealing with biochemical induction in vitro: different molecules must be supplemented according to a specific time-program in order for the treatment to be effective. The material features discussed so far, able to affect cell fate and functions, are inherently static in nature. Examples of dynamic changing material features have already been reported. For example, materials that vary adhesiveness and stiffness in a user-defined manner proved to be effective in changing cell behavior dynamically.71,72 An issue may arise when applying the stimulus for changing the feature of interest: it must not alter cell state or affect viability. In the two previously cited works, the triggers were mechanical stretching of the substrate - which reversibly exposes/hides functional sites, and UV irradiation respectively. In both examples, cells responded to the trigger promptly, within a few hours, and changed their behavior accordingly. Topographic changes also proved to dynamically affect cell ACS Paragon Plus Environment

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adhesion and orientation, as recently reported by Zhernenkov et al. by means of a patterned thermoresponsive polymer.73 Along this line, we have recently employed light sensitive azopolymers to create dynamic micronscale topographic patterns.74 Gratings in the form of parallel channels or square pillars were produced by means of structured light generated by a Lloyd’s mirror setup. NIH3T3 fibroblasts adhered and recognized the patterned signals without further surface functionalization. The pattern was then manipulated by means of incoherent light in presence of attached living cells. The use of incoherent light did not affect cell viability, demonstrating the feasibility of the process. The exploitation of time-changing signals according to pathways defined ab initio is still in its infancy. Yet, we strongly believe that the introduction of the “time-dimension” as control parameter for cell response induction would enormously amplify the governing power of material surfaces. This would be beneficial in those applications requiring a fine control over cell behavior, such as design of scaffolds for tissue engineering, chips for drug/molecules screening in vitro or active devices for cell therapy.

CONCLUSION Biochemical and biophysical signals displayed by material surfaces proved to profoundly affect cell behavior. Cell adhesion complexes and cytoskeletal generated forces play a significant role in transducing these signals into genetic events that eventually dictate cell fate and functions. Recent advancements in micro- and nano-fabrication technologies enabled a tight control on the spatial positioning and presentation of material cues. This allowed to dissect cell recognition of and response to patterns of biochemical, topographic and mechanical stimuli. Although the molecular mechanisms governing the process of cell-signal interaction have not yet been thoroughly understood, exploiting material stimuli to control cell behavior can potentially be employed in several biotechnological and clinical settings. In fact, materials able to control cell differentiation and morphogenetic events can be used to create systems to supplement and screen drugs in microenvironments that more closely mimic ACS Paragon Plus Environment

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native ones, or to engineer novel prosthesis or devices that better interact with the surrounding biological host, thus improving their performances in situ.

ASSOCIATED CONTENT Supporting Information Experimental details and image acquisition methods. This material is available free of charge via the Internet at http://pubs.acs.org.

AUTHOR INFORMATION Corresponding Author *E-mail: [email protected] Author Contributions The manuscript was written through contributions of all authors. All authors have given approval to the final version of the manuscript. Notes The authors declare no competing financial interest.

ACKNOWLEDGMENT The authors are grateful to acknowledge Prof. F. Causa, Dr. M. Iannone, Ms. V. La Tilla and Dr. C. F. Natale for providing original materials for the assembling of figure panels. The authors thank Mrs. R. Infranca for the precious assistance during proofreading.


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