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Magnetically tuning tether mobility of integrin ligand regulates adhesion, spreading, and differentiation of stem cells Siu Hong Dexter Wong, Jinming Li, Xiaohui Yan, Ben Wang, Rui Li, Li Zhang, and Liming Bian Nano Lett., Just Accepted Manuscript • DOI: 10.1021/acs.nanolett.6b04958 • Publication Date (Web): 24 Feb 2017 Downloaded from http://pubs.acs.org on February 27, 2017
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Magnetically Tuning Tether Mobility of Integrin Ligand Regulates Adhesion, Spreading, and Differentiation of Stem Cells Dexter S. H. Wong#,
|| ,
Jinming Li#, ||, Xiaohui Yan||, Ben Wang||, Rui Li||, Li Zhang|| and
Liming Bian*, ||, ‡, ¶, §, ζ ||
Department of Mechanical and Automation Engineering (Biomedical Engineering), ‡Shun
Hing Institute of Advanced Engineering, The Chinese University of Hong Kong, Hong Kong, China, ¶Shenzhen Research Institute, The Chinese University of Hong Kong, Hong Kong, China, §China Orthopedic Regenerative Medicine Group (CORMed), Hangzhou, China, ζ
Centre for Novel Biomaterials, The Chinese University of Hong Kong, Hong Kong, China
# These authors contributed equally to this work *To whom correspondence should be addressed. E-mail:
[email protected] Graphical Table of Contents
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Abstract: Cells sense and respond to the surrounding microenvironment through binding of membranous integrin to ligands such as the Arg-Gly-Asp (RGD) peptide. Previous studies show that the RGD tether properties on substrate influence cell adhesion and spreading, but few studies have reported strategies to control the tether mobility of RGD on substrate via a physical and non-contact approach. Herein, we demonstrate a novel strategy to tune the tether mobility of RGD on substrate via magnetic force. We conjugate a monolayer of RGD-bearing magnetic nanoparticles (MNPs) on glass substrate via the flexible and coiled poly(ethylene glycol) linker of large molecular weight (PEG, average MW: 2000), and this increases the RGD tether mobility, which can be significantly reduced by applying magnetic attraction on MNPs. Our data show that high RGD tether mobility delays the early adhesion and spreading of human mesenchymal stem cells (hMSCs), leading to compromised osteogenic differentiation at later stage. In contrast, hMSCs cultured on substrate with restricted RGD tether mobility, achieved either via a shorter PEG linker (MW: 200) or magnetic force, show significantly better adhesion, spreading, and osteogenic differentiation. The control utilizing RGD-bearing non-magnetic nanoparticles shows no such enhancing effect of magnetic field on cellular events, further supporting our conjecture of magnetic tuning of RGD tether mobility. We hypothesize that high tether mobility of RGD entails additional time and effort by the cells to fully develop traction force and mechanical feedback, thereby delaying the maturation of FAs and activation of subsequent mechanotransduction signaling. Our staining results of vinculin, a critical component of FAs, and Yes-associated protein (YAP), an important mechanosensitive transcriptional factor, support our hypothesis. We believe that our work not only sheds light on the impact of dynamic presentation of cell adhesive ligands on cellular behaviors, which should be taken into consideration for designing novel
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biomaterials, but also formulate an effective non-contact strategy that enables further investigation on the mechanobiological mechanisms underlying such cellular responses. Keywords: tether mobility, magnetic nanoparticles, human mesenchymal stem cells, cellsubstratum interactions, stem cell differentiation
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Cells, especially human mesenchymal stem cells (hMSCs) have the ability to sense and respond to the physical nature of their microenvironment, particularly the components of the extracellular matrix (ECM).1-3 Cell adhesion to the surrounding ECM is important to many cellular functions4-7 and is mediated primarily by the integrin family, a group of heterodimeric transmembrane proteins.8-10 Integrin-mediated formation of focal adhesions (FAs) plays a critical role in cellular adhesion and sensing of the external environment and the subsequent initiation of cascading intracellular signaling through a process known as mechanotransduction.11 Surface conjugation of cell adhesive motifs including ECM molecules or ECMderived bioactive peptides (e.g., Arg-Gly-Asp/RGD peptide) is critical to the development of cell traction force and subsequent cell adhesion, spreading, and mechanosensing on non-cell adhesive substrates.12,
13
As the anchorage of the cell adhesion, the physicochemical
presentation of these conjugated cell adhesive molecules has substantial impact on their capacity to mediate traction force development and the aforementioned cellular behaviors. For instance, previous studies have shown that the spacing between the conjugated RGD peptide significantly influences the integrin clustering during the formation of focal adhesions and subsequent the adhesion and spreading of cells on the substrate.14 Furthermore, the viscoelastic properties of the 2D substrate and 3D hydrogel scaffold that are grafted with cell adhesive ligands significantly influence the spreading and differentiation of the seeded hMSCs.15,
16
Another recent study also suggested that the tether density of fibrous ECM
molecules including fibronectin and collagen on substrate regulates cellular sensing of the substrate rigidity, and loosely bound and flexible ECM molecules reduce the mechanical feedback by cells via traction force development.17 Similarly, one report revealed that short tether length is equivalent to stiff matrix while long tether length is equivalent to soft matrix, which combine as a matrix interplay factor to influence the cell mechanosensing signaling.18 4 ACS Paragon Plus Environment
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In addition, a study also showed that cells only develop mature FAs on substrate coated with a lipid layer containing RGD peptide of restricted mobility, and fully mobile RGD leads to the formation of podosome like adhesion in the absence of traction force.19 Moreover, we previously reported that RGD peptide coupled to the substrate by stronger electrostatic force better promote the adhesion, spreading, and differentiation of hMSCs than the RGD coupled by weaker force.20 These findings indicate that the molecular compliance (deformation per unit force) of the tethered cell adhesive ligands may play an important role in regulating traction force development, cell adhesion and mechanosensing. However, little research attention has so far been dedicated to this topic. Modulation on the physicochemical presentation of cell adhesive ligands enables dynamic control of cell-substrate interactions, which is of great significance to fundamental researches and practical applications.21 Different strategies, including host-guest chemistry22, 23
, photochemistry24,
25
, electrochemistry26, thermal induction27 and enzymatic cleavage28,
have been reported to alter the presentation of cell-adhesive peptide on substrate to regulate the cell adhesion or detachment. Ultimately, most of these strategies involve chemical intervention to alter the tethering of RGD peptide and may interfere with normal cellular events, and the resulting compounding factors hamper elucidation of cellular response to substrate dynamics. A direct, physical, and non-contact strategy to modulate the tethered cell adhesive ligands and control the cell adhesion behavior on substrate will be highly desirable to minimize any potential interference on cells. Herein, we describe a simple and facile technique to tune the mobility of RGD peptides tethered on substrates by using magnetic field as graphically illustrated in Figure 1. Briefly, we conjugated a monolayer of RGD-grafted magnetic nanoparticles (Fe3O4,, MNPs coated with silica) via a poly(ethylene glycol) linker (PEG, average molecular weight: 2000 to glass substrate, and blocking agent (bovine serum albumin, BSA) was coated to the 5 ACS Paragon Plus Environment
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substrate to reduce non-specific cell adhesion before seeding of hMSCs.29 Our finding revealed that a downward magnetic attraction to the tethered RGD-bearing MNPs significantly improved the initial cell attachment and subsequent spreading on the substrate until up to day 7 compared to the control without exposure to the magnetic field. Replacing the MNPs with non-magnetic pure silica nanoparticles (SiO2 NPs) abrogated such enhanced cell adhesion and spreading under magnetic field. Table 1 lists all the group name abbreviations including experimental groups, control groups, positive and negative control groups which will be used in the later text. The group is applied with magnetic field is denoted as +Mag at the end of the abbreviation while no applied magnetic field is denoted as -Mag. Furthermore, we demonstrated that the hMSCs seeded on magnetic responsive substrate exhibit enhanced activation of mechanotransduction signalling involving the Yesassociated protein (YAP) transcriptional regulator and osteogenic differentiation compared to the controls. We speculate that the presence of the lengthy and flexible PEG linker between glass substrate and RGD bearing MNPs engenders additional nanoscale tether mobility and compliance (stretchability) of RGD that may retard the development of cell traction force and therefore cell adhesion and spreading. Application of magnetic field to pull the RGD bearing MNPs firmly against glass substrate may substantially reduce this tether compliance, thereby facilitate the traction force development and adhesion, spreading, and mechanotransduction mediated osteogenic differentiation of hMSCs. To the best of our knowledge, our work is the first to demonstrate that altering tethering mobility of RGD peptides on substrate via magnetic field regulates the adhesion, spreading, mechanosensing and differentiation of hMSCs. The technique developed in this study to control cell-substrate interactions via noncontact physical mechanism will also be valuable to advances in fundamental researches. We first synthesize MNPs of ~ 20 nm in diameter (Supplementary Fig. S1a), and amino-functionalized silica coated MNPs (MNP@SiO2) (Supplementary Fig. S1b) and 6 ACS Paragon Plus Environment
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amino-functionalized silica nanoparticles (SiO2) NPs (Supplementary Fig. S2) of ~ 40 nm in diameter as previously reported.30, 31 Transmission electron microscopy (TEM) images reveal that the physical sizes of synthesized MNPs, MNP@SiO2 and SiO2 NPs are 15.9 ± 2.7 nm, 36.6 ± 2.4 nm and 38.3 ± 2.4 nm respectively. Dynamic light scattering (DLS) measurement shows that the hydrodynamic size of MNP, MNP@SiO2 and SiO2–NH2 NPs are 18.5 ± 1.2 nm, 43.6 ± 3.4 nm, and 43.8 ± 2.6 nm respectively (Supplementary Fig. S3a, b and d), which are consistent with the TEM data. Zeta potential of MNP, MNP@SiO2 and SiO2NPs reveal their surface charges of –OH and –NH2 in deionised water are -9.9 ± 5.2 mV, 27.9 ± 7.1 mV and 31.5 ± 3.9 mV respectively (Supplementary Fig. S3f, g and i).32 Vibrating Sample Magnetometer (VSM) shows that the magnetic moment per unit mass of MNP and MNP@SiO2 NPs exhibit minimal hysteresis, and their saturation moments (Msat) are estimated to be 54 emu g-1 and 28 emu g-1, respectively (Supplementary Fig. S1c). The distance between the coverslip surface (and hence the MNPs) and the permanent magnet is approximately 1 mm. The exerted magnetic strength on MNPs is estimated to be 86.1 ± 1.01 mT (Supplementary Fig. S1d), and we believe the resulting magnetic attraction force reduces the mobility and tether compliance of the RGD-grafted MNPs on the glass substrate. Next, we PEGylate the MNP@SiO2 and SiO2 NPs with PEG (MW = 2000) to form the MNPs@SiO2–PEG2000 and SiO2–PEG2000 with a hydrodynamic size of 61.2 ± 3.6 nm and 62 ± 4.7 nm, and this increase in the hydrodynamic diameter is similar to the reported length of the PEG linker in the literature.33 A significant drop of zeta potential from 32.7 ± 3.4 mV to 0.81 ± 6.8 mV for (Supplementary Fig. S3h) MNPs@SiO2–PEG2000 and from 31.5 ± 3.9 mV to 8.2 ± 5.3 mV for (Supplementary Fig. S3j) SiO2–PEG2000 probably suggests the successful conjugation of the PEG linker which has a relatively neutral zeta potential.34 We subsequently characterize the particles by Fourier transform infrared (FTIR) spectroscopy to confirm the successful silica coating and conjugation of RGD (Supplementary Fig. 1e). The
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FTIR spectrum of MNP (black line) shows a band at 576 cm-1, which is attributed to Fe–O stretching.35 At the same spectrum, O–H deformed vibration near 1620 cm-1 and FeO–H stretching vibration near 3410 cm-1 were observed. The FTIR spectrum of MNPs@SiO2 (red line) exhibits the peak at 876 cm-1 and 1000 cm-1 for Si–O and Si–O–Si stretching respectively, which indicate the presence of silica coating.36 The new peaks at around 2850-2950 cm-1 and broad peak around 1200 cm-1 are associated with the C–H stretching and the C–O–C stretching vibration in the PEG linker of MNPs@SiO2–PEG2000 (deep blue line).35 The strong peak at 1705 cm-1 and the shoulder at 1650 cm-1 contribute to the C=O amide and the N–H bending vibration from the maleimide groups of the PEG linker.37 After the grafting of RGD peptide, the spectrum of MNPs@SiO2–PEG2000–RGD (pale blue line) exhibits a new band at 1377 cm-1 and 1421 cm-1 for stretching vibration and deformation vibrations of amide III bond (–CO–NH–).38 Superparamagnetic iron oxide nanoparticles (SPIOs) or magnetic nanoparticles (MNPs), have been clinically approved for non-invasive diagnose of reticuloendothelial diseases.39 It has also been shown that the cellular uptake (9.6 ± 0.3 pg, approximately 1.48 million MNP per cells) of PEG-coated MNPs has insignificant cytotoxicity and influence on the differentiations (osteogenesis, chondrogenesis and adipogenesis) of hMSCs for up to 7 days of culture.40 Furthermore, silica coated MNPs have been shown to be highly stable in culture medium (no detectable Fe ions in culture medium at 37 oC for up to 96 h),41 and the further passivation by PEG conjugation in our work provides extra colloidal stability of MNPs. These data suggest that the MNPs used in our study are suitable vehicles for presenting cell adhesive ligands. We fabricate a monolayer of RGD bearing MNPs substrate for tunable tether mobility. After each step of surface coating, water contact angle measurement is firstly performed to confirm the successful modifications. Before silanization, piranha is used as a typical strong oxidizing agent to activate the hydroxyl groups of glass surface.42 The water contact angle 8 ACS Paragon Plus Environment
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measured on the piranha treated glass decreases from (Fig. 2ai) 72.5 ± 2o to (Fig. 2aii) 46.7 ± 2o, and this shows the increased substrate hydrophilicity due to an increased number of hydroxyl groups. The substrate contact angle increases to (Fig. 2aiii) 69 ± 1.3o after MPTMS silanization, which is close to the previously reported value of highly organized MPTMS selfassembled monolayers (SAMs) with surface –SH groups.43 The substrate becomes hydrophilic again with reduced contact angle (Fig. 2aiv) 41.3 ± 1.5o after grafting MNPs to the surface because of the hydrophilic PEG linker from the grafted MNPs surface.44 To further confirm the successful fabrication, scanning electron microscopy (SEM) imaging and atom force microscopy (AFM) imaging show a uniform monolayer of MNPs grafted on the glass surface (Fig. 3b and Fig. S4). The density of SiO22000, MNP200 and MNP2000 on substrate measured by SEM are 17 ± 2, 15 ±1 and 18 ± 3 particles/µm2 respectively (Fig. 2b); measured by AFM imaging are 18 ± 3, 14 ± 2 and 15 ±3 particles/µm2 respectively (Supplementary Fig. S4), which are consistent to each other. The spacing and density of RGD have been shown to affect and regulate stem cell behaviours.4, 45, 46 Therefore, the density of grafted MNPs is kept relatively constant for all groups by using the same fabrication method. In addition, the average RGD density per particle on MNP200, MNP2000 and SiO22000 are 1767, 1683 and 1752 peptides/particle (0.42, 0.37 and 0.4 peptide/nm2) respectively, which have a relatively constant cluster of RGD peptides on a particle. Next, we demonstrate the restriction on the tether mobility of RGD bearing MNPs by an external magnetic field (Fig. 2ci). We use AFM contact mode to apply a swiping force of increasing magnitude on the MNPs by increasing the deflection setpoint voltage (V) from 1.000 V to 10.000 V. We observe that the MNPs are almost cleared by the AFM tip when setpoint voltage is increased to 10.000 V in the absence of magnetic field (Fig. 2cii). In contrast, less particles are swiped away from the substrate at the same voltage in the presence of magnetic field (magnetic field strength at ~40 mT and ~80 mT in Fig. 2ciii and Fig. 2civ, 9 ACS Paragon Plus Environment
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respectively). The result suggests that the magnetic attraction force significantly restricts the mobility of the RGD bearing MNPs around the substrate tether. Based on this developed dynamic substrate with tunable tether mobility of RGD peptide via magnetic field, we further investigate the effect of such varying tether mobility of RGD on stem cell adhesion, spreading and differentiation. To study the early response of hMSCs to the tether mobility of RGD, we seed hMSCs on the glass substrate modified with RGD-bearing MNPs with and without continuous downward magnetic field for 24 h in basal medium before examining cell adhesion and cell spreading. From Table 1, as a parallel control, non-magnetic silica nanoparticles are used to replace the MNPs to bear RGD peptides (SiO22000R). In addition, a negative control employing bare MNPs without RGD (MNP2000) is also included, and a positive control group utilizes a significantly shorter PEG linker (MW: 200) for the conjugation of MNPs (MNP200) to substrate to produce RGD-bearing MNP (MNP200R) with intrinsically low tether mobility. Strikingly, the different substrate surfaces show substantially different amount of cell adhesion after 24 h (Fig. 3). The substrate conjugated with MNP2000+/-Mag (RGD absence) show few adherent cells with little spreading, indicating the necessity of RGD for cell adhesion and spreading observed by fluorescence micrographs (Fig. 3 and 4). More interestingly, the number of adherent and spreading cells in the immobilized RGDbearing MNP200R+/-Mag and MNP2000R+Mag are significantly higher than that in the mobile RGD-bearing surface SiO22000R+/-Mag, and MNP2000R-Mag, and no statistical significance are found within the mobile RGD-bearing surfaces and within the immobilized RGD-bearing surfaces (Fig. 3c). For example, MNP2000R+Mag group has 207% more adherent cells than that of MNP2000R-Mag group (immobilized vs mobile RGD-bearing surfaces). And SiO22000R has 155% more adherent cells than that of MNP2000 group (RGD-bearing vs RGD-absence surfaces). Hence, we draw a trendline to shortly conclude 10 ACS Paragon Plus Environment
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that tether mobility (black line) is inversely related to initial cell adhesion (red line) (Fig. 3a & b). In addition, we have included the negative control MNP2000 without RGD peptide conjugation. It has been shown to have minimal effect on cell adhesion and cell spreading for hMSCs (Fig. 3a & b). To confirm the relationship between RGD tether mobility and cell adhesion, we evaluate the cell morphology, spreading area and vinculin expression, which is an important adaptor protein of FAs (Focal adhesions), after 24 h of cell culture (Fig. 4). We observe that cells grown on substrates with high RGD tether mobility (SiO22000R+/-Mag and MNP2000R-Mag) develop a more spindle-like and elongated morphology with significantly larger cell shape factor (i.e., aspect ratio, major/minor axis) but smaller cell size (Fig. 4a). On the other hand, hMSCs are significantly more spread on the substrates with restricted RGD mobility (MNP200R+/-Mag and MNP2000R+Mag) with smaller shape factor but larger cell size (Fig. 4b). For example, the cell spreading area measured MNP2000R+Mag has 161% higher than that of MNP2000R-Mag (Fig. 4c). Similarly, the shape factor of cells cultured on immobilized RGD is close to unity and significantly lower than that of the cells cultured on RGD of higher tether mobility as the result of more substantial spreading in all directions (Fig. 4d). Furthermore, hMSCs grown on immobilized RGD develop more pronounced actin stress fibres and punctate vinculin accumulation at sites of FAs, which mediate cell adhesion and mechanotransduction. In contrast, the filamentous actin development and punctate vinculin accumulation is substantially less evident in hMSCs cultured on substrates with high RGD tether mobility. In addition, for the surface without RGD conjugation, cells spread poorly, and vinculin staining is hardly found. Our finding suggests that restricted RGD tether mobility on the substrate favours the maturation of FAs and cell spreading, while high RGD tether mobility hinders the formation of FAs and cell spreading. We further culture the hMSCs on these four groups of substrates in basal medium for 11 ACS Paragon Plus Environment
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an additional 2 days, and the differential adhesion and spreading of hMSCs in different groups are still evident after 3 days of culture (Fig. 5). Generally, cells in all groups spread more and develop better focal adhesions on day 3 compared to day 1 (Fig. 4 & 5). This agrees with the previous study that cell adhesion and spreading on substrate is time dependent.47, 48 Cells grown on MNP2000+/-Mag substrate still show the least cell spreading among all the groups. The cell size of the groups with high RGD tether mobility (SiO22000R+/-Mag and MNP2000R-Mag)
or
restricted
RGD
tether
mobility
(MNP200R+/-Mag
and
MNP2000R+Mag) increases by around 60% or 80% respectively from day 1 to day 3. However, it is worthy to note that hMSCs cultured on substrate with restricted RGD mobility still show around 100% larger cell size and 50% lower shape factor compared to that of the groups with mobile RGD (Fig. 5c & d). Furthermore, there is still significantly more punctate vinculin staining in the groups of restricted RGD mobility than in the groups of mobile RGD (Fig. 5a to b). It is known that the formation and maturation of FAs requires active actomyosinmediated mechanical feedback between the extracellular environment and intracellular cytoskeletal structures upon the initial ligation of integrin receptor and its ligands presented on the substrate. We believe that the PEG linker used for the conjugation of RGD-bearing MNPs may have a critical impact on the formation and maturation of FAs by cells in our study. PEG molecules, especially those of high molecular weights, are known to have a highly flexible and coiled chain49-51 that afford the inherent mobility and compliance to the tethered RGD-bearing MNPs. A previous study actually showed that cells can stretch and unfold the coiled PEG linker by attaching to and pulling the conjugated RGD, and this enables estimation of the cell traction force based on the extent of PEG unfolding.52 In our study, upon RGD ligation with the integrin on hMSCs upon initial cell adhesion, we speculate that this inherent tether compliance of RGD due to the PEG linker entails additional 12 ACS Paragon Plus Environment
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time and effort by the cells to stretch the coiled PEG linker, and this delays the development of cell traction force, which is important to the maturation of FAs.53 This is consistent with another previous study which showed that presentation of mobile RGD based on polyrotaxane threading also delays the formation of mature FAs and spreading of cells on the substrate despite faster cell attachment within the first ~20 minutes of cell seeding.54 On the other hand, the shorter PEG linker (MW: 200) possesses significantly less chain flexibility and coiling and therefore has less influence on the adhesion and spreading of cells. This speculation is supported by the substantial cell adhesion and spreading regardless of the application magnetic field when the shorter PEG linker is used (MNPR200+/-Mag) (Fig. 3 to 5). Furthermore, one powerful advantage of our unique design is that when a magnetic field is applied, the magnetic attraction firmly anchors the RGD bearing MNPs on the substrate and effectively restrict the tether mobility of RGD by overriding the intrinsic mobility and compliance of the PEG linker. Our data indeed show that high RGD mobility due to the longer PEG linker (MW: 2000) may hinder the mechanical sensing and feedback of cells via integrin ligation, thereby leading to the limited cell adhesion and spreading on such substrate. In contrast, the restricted RGD tether mobility either due to shorter PEG linker (MNP200R+/Mag) or magnetic attraction (MNP2000R+Mag) on the RGD-bearing MNPs likely facilitate the mechanosensing of the cells upon integrin ligation with RGD, thereby promoting the early adhesion and spreading of cells.17 We postulate that it takes longer time for the cells on the substrate with high RGD tether mobility to fully unfold the PEG linker before developing mature FAs and spreading further. Although each MNP provides multiple strands of RGD peptide in a cluster form, which is important to the clustering of integrins,18 the length/mobility of the tether between MNP and substrate still dictates the time needed for the cells to form mature FAs via mechanical feedback. It is noted that cells start to spread in spindle-like shape with increasing size after 3 days on the non-fouling control surface
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(MNP2000) which is blocked with BSA protein. This phenomenon is consistent with the previously reported finding that the blocking effect of BSA coating diminishes over time.29, 55 Nevertheless, such non-specific cell spreading is still very small when compared to the RGDmediated cell spreading in other groups. Furthermore, we observed that MNP200R-/+Mag groups have slightly better cell adhesion (Fig. 3c), day 3 cell spreading (Fig. 5c), cell differentiation (Fig. 6b & 7b) and YAP nuclear localization (Fig. 8b) than the MNP2000R+Mag group, but the difference is not statistically significant. Hence, we didn’t further compare results between these groups. It has been shown that the tether length of cell adhesive ligands affects the size and length of focal adhesions and influences cell spreading and attachment.18 By using various tether length (molecular weight) of Mal-PEG-NHS ester, Attwood et. al have demonstrated that the tether length of cell adhesive ligands at around 40 nm constitutes a critical threshold of cellular sensing distance above which cells cannot effectively develop the critical traction force (>43 pN critical force to activate integrins) that is required for developing mature focal adhesions and cell spreading. In our work, the tether of RGD to the glass consists of the ~17.8 nm initial PEG spacer between MNP and substrate56, the intermediate MNP of ~36.6 nm in diameter, and the ~17.8 nm PEG spacer between MNP and RGD. Therefore, the total tether length is ~72.2 nm (including the diameter of MNP). In the absence of magnetic attraction, the longer RGD tether length entails extra time and effort by the cells to develop the critical traction force required for efficient cell adhesion and spreading. On the other hand, magnetic attraction instantly decreases the tether length from ~72.2 nm to ~17.8 nm, and this represents an almost 4 times reduction in the RGD tether length. Besides, this shortened tether length is within the range of the critical tether length (
