Incorporation of F-MWCNTs into electrospun nanofibers regulates osteogenesis through stiffness and nanotopography

Materials Science & Engineering C 106 (2020) 110163

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Incorporation of F-MWCNTs into electrospun nanofibers regulates osteogenesis through stiffness and nanotopography ⁎

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Fatemeh Jahanmarda,b,c, , Mohamadreza Baghban Eslaminejadb, , Mohammad Amani-Tehrand, Fatemeh Zareib, Naeimeh Rezaeie, Michiel Croesa, Saber Amin Yavaria a

Department of Orthopedics, University Medical Centre Utrecht, Heidelberglaan 100, 3584, CX, Utrecht, the Netherlands Department of Stem Cells and Developmental Biology, Cell Science Research Center, Royan Institute for Stem Cell Biology and Technology, ACECR, P.O. Box: 16635-148, Tehran, Iran c Nanotechnology Institute, Amirkabir University of Technology, P.O. Box: 15875-4413, Tehran, Iran d Department of Textile Engineering, Amirkabir University of Technology, P.O. Box: 15875-4413, Tehran, Iran e Department of Cell and Molecular Biology, Cell Science Research Center, Royan Institute for Biotechnology, ACECR, Isfahan, Iran b

A R T I C LE I N FO

A B S T R A C T

Keywords: Stiffness Nanoroughness Osteogenic differentiation Bone substitute Electrospinning

Nanotopography and stiffness are major physical cues affecting cell fate. However, the current nanofiber modifications techniques are limited by their ability to control these two physical cues irrespective of each other without changing the materials' surface chemistry. For this reason, the isolated effects of topography and stiffness on osteogenic regulation in electrospun nanofibers have been studied incompletely. Here, we investigated 1. how functionalized multiwall carbon nanotubes (F-MWCNTs) loaded in Polycaprolactone (PCL) nanofibers control their physical properties and 2. whether the resulting unique structures lead to distinctive phenotypes in bone progenitor cells. Changes in material properties were measured by high-resolution electron microscopes, protein adsorption and tensile tests. The effect of the developed structures on human mesenchymal stem cell (MSC) osteogenic differentiation was determined by extensive quantification of early and late osteogenic marker genes. It was found that F-MWCNT loading was an effective method to independently control the PCL nanofiber surface nanoroughness or stiffness, depending on the applied F-MWCNT concentration. Collectively, this suggests that stiffness and topography activate distinct osteogenic signaling pathway. The current strategy can help our further understanding of the mechano-biological responses in osteoprogenitor cells, which could ultimately lead to improved design of bone substitute biomaterials.

1. Introduction Synthetic bone substitutes have received tremendous attention to replace natural tissue and treat bony defects, considering the complications and limited tissue supply associated with autologous bone grafting [1]. Biomimetic scaffolds that provide the native physical, chemical and biological cues for osteoprogenitor cells are thought to better recapitulate physiological bone healing [2]. In particular, nanofibrous biomaterials are considered to be effective structures for tissue regeneration because they typically mimic the native extracellular matrix (ECM) structure [3,4] to enhance cell spreading [5], adhesion [6,7], and differentiation [8]. In fact, the differentiation of mesenchymal stem cells (MSCs) into tissue-specific cells is for a large part orchestrated by the physical signals provided by the ECM

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microenvironment [9,10]. More specifically, it has been proven that the nanotopography and stiffness are among the most crucial factors of the biomaterial governing cell differentiation [11–14], including the osteogenic behaviour of MSCs [12,15]. Many studies have demonstrated the importance of physical signaling in stem cell differentiation in different types of biomaterials [16,17]. However, its role in electrospun ECM-like nanofibers has only recently received attention. A number of studies indicate that an increase in surface nanoroughness and stiffness of electrospun scaffolds is beneficial for the osteogenic response [18–20]. Nevertheless, the isolated effects of these two physical cues in electrospun scaffolds remain elusive. Different techniques have been applied to explore the mechano-biological effects of electrospun nanofibers, but can only verify one of the physical cues, i.e. fiber annealing [18], UV crosslinking [21],

Correspondence to: F. Jahanmard, Department of Orthopedics, University Medical Centre Utrecht, Heidelberglaan 100, 3584 CX, Utrecht, the Netherlands. Corresponding author. E-mail addresses: [email protected] (F. Jahanmard), [email protected] (M. Baghban Eslaminejad).

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https://doi.org/10.1016/j.msec.2019.110163 Received 9 July 2019; Received in revised form 16 August 2019; Accepted 4 September 2019 Available online 05 September 2019 0928-4931/ © 2019 Elsevier B.V. All rights reserved.

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scanning electron microscope (SEM) (Seron Technologies AIS2100, Korea) and coating with a thin layer of gold. The distribution and average fiber diameter were calculated for 50 nanofibers in each image using ImageJ (version 1.4.3.67, National Institutes of Health, US) and repeated for a total of five images per group. The porosity of the electrospun nanofibers were calculated using the Eq. (1) [19]:

and use of core-shell structures [20] to produce stiffer nanofibers, and applying surface porosity [22,23], nanoimprint lithography [24], physical coatings [25] to produce nanoroughness. Therefore, there is need of a single method that can be applied to optimize both surface nanoroughness and stiffness, the two crucial parameters involved in regulating cell function, in the same biomaterials satisfactorily. Moreover, for an optimal comparison of roughness or stiffness-induced responses, the chosen modification method should not interfere with the surface chemical properties, since this could mask the true mechano-biological responses. In light of the aforementioned challenges, it has been found that the incorporation of carbon nanotubes (CNTs) into electrospun nanofibers is a relatively straightforward means to modulate their stiffness and surface nanoroughness [26,27]. CNTs role to improve mechanical, physical and electrical properties of matrix in other advanced materials have been also acknowledged [28,29]. Importantly, the orientation of CNTs plays a crucial role in determining the nanofiber physical properties, which can be controlled by using the appropriate CNT concentration [30–32]. The homogenous dispersion of a relatively low CNT concentration in the polymer matrix supports its natural alignment in the nanofibers, leading to enhanced material stiffness [33–37]. Contrarily, a relatively high CNT concentration leads to more heterogeneous dispersion and formation of CNT bundles followed by increasing nanoroughness of the nanofiber surfaces [38]. Despite this evidence, there has been no comprehensive investigation of the effect of CNT incorporation, and its concentration-dependency, on the physical and pro-osteogenic properties of electrospun nanofibers. The aim of this study was to determine if the incorporation of CNTs into nanofibers can modify the topographical and mechanical properties of nanofiber structures irrespective of each other, and subsequently verify the role of these properties on osteogenic differentiation. To this end, we fabricated nanocomposite structures composed of PCL nanofibers with functionalized multiwall carbon nanotubes (F-MWCNTs). The scaffold compositions were optimized to maximize the stiffness and nanoroughness of PCL nanofibers without changing the surface chemistry. To determine how PCL/F-MWCNTs scaffolds with unique physical properties alter osteogenic regulation differently, we evaluated their modulatory effect on a panel of osteogenic marker expression.

ρ P = ⎜⎛1 − ⎟⎞ × 100 ρ′ ⎠ ⎝

(1)

where P is the porosity of the matrix, ρ is the apparent density of matrix and the ρ′ is the density of polymer materials used to fabricate the nanofibers. To analyze the distribution and orientation of the F-MWCNTs in the PCL matrix, transmission electron microscopy (TEM) (FEI Tecnai 12, Thermosystems-FEI, Netherlands) was used. The nanofibers were prepared on Formvar-coated grids with a carbon layer. Then, the grids were evaluated on a TEM with a digital camera. To characterize the fiber nanoroughness, atomic force microscopy (AFM, Ambios tech, US) was employed with a Nanosensors probe (silicon cantilever) under tapping mode in an air atmosphere. For each matrix, a scan of 20 × 20 mm was made for a precise test (n = 5). The average roughness (Ra) was measured using Scanning probe image processor (SPIP) software (version 6.7, Image Metrology A/S, Denmark). Prior to measuring the nanoroughness via the AFM images, single nanofibers were flattened to estimate the average roughness on the nanofiber surfaces. FTIR spectra were obtained using a FTIR Spectrometer (Thermo Fisher Scientific, Nicolet iS10, US) in the attenuated total reflection mode. 2.2.1. Protein adsorption assay Bovine serum albumin (BSA) was selected as a reference protein to evaluate protein adsorption to the nanofiber surface, based on a previous report [39]. The matrices were cut into 10 × 10 mm squares, soaked in methanol for 15 min, and rinsed three times with phosphate buffered saline (PBS). Next, they were covered in PBS in a 24-well plate and incubated at 37 °C overnight to pre-wet them. The wetted nanofibers were immersed in a BSA solution (4 mg/ml) at 37 °C overnight. The samples were transferred to a new tube and washed three times with PBS to remove non-adherent proteins. The amount of adsorbed protein was measured using a bicinchoninic acid (BCA) protein assay kit (Merck, Novagen, Germany) according to the manufacturer's instructions. The amount of protein was normalized to the surface area for each sample [40], as shown in Eq. (2):

2. Materials and methods 2.1. Preparation of PCL/F-MWCNT composite nanofibers A suspension of Poly (ε-Caprolactone) PCL (MW = 80,000; Sigma, Germany) was prepared by dissolving 11 wt% PCL in chloroform. Various amounts of COOH-Multiwall carbon nanotubes (FMWCNTs, > 95%, OD: 20–30 nm, US Research Nanomaterials, US) were dispersed in methanol and sonicated by a probe sonicator (Hielscher Ultrasound Technology UIP1000hd, Germany) for 1 h to achieve a homogeneous dispersion. To make the PCL/F-MWCNT solution, the F-MWCNT solutions were added dropwise to the PCL solutions in an F-MWCNT:PCL volume ratio of 1:3 while stirring. The final concentrations of the F-MWCNTs were 0, 0.5, 1, 2, and 3 wt% with reference to the PCL mass in the solution. The PCL/F-MWCNT solution was sonicated in an ultrasonic bath (Elma, Transsonic 460, Germany) for 1 h and electrospun in a 1 ml plastic syringe at a flow rate of 1 ml/h. The distance between the needle tip and the collector was adjusted to 17 cm and the solution was stretched under the high voltage of 18 kV. This procedure was performed under controlled conditions of 35 °C and 35% relative humidity. The obtained nanofibers were collected on a grounded rotary drum. Finally, the nanofibers were dried under vacuum to remove any residual solvents.

SD =

4m ρD

(2)

where SD is the surface density of the nanofibers, m is the mass of the nanofibers, ρ is the intrinsic density of the polymer and D is the average diameter of the nanofibers. 2.3. Mechanical characterization The mechanical properties of the specimens were measured by a 100 N Instron (5566 universal testing machine, US) load cell, using crosshead rates of 10 mm/min (5 × 20 mm nanofiber matrix) and 1 mm/min (20 x 20 mm aligned nanofibers). In order to determine the stiffness of a single nanofiber, we used a model that was developed in our previous work [41]. Briefly, the young modulus of single nanofibers was calculated based on the Eq. (3) in which the modulus of aligned nanofibers was already measured by the tensile test.

ET 1 = Es γπ

2.2. Characterization The morphology of the electrospun nanofibers was observed by a

π 2

∫− π 2

1 cos θ dθ (1 + ((θ / γ )2)

(3)

where ETis the Young's modulus of an aligned nanofiber, ES is the 2

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Young's modulus of single nanofiber, γ and θ refer to the standard deviation and the orientation angle of the nanofibers in the matrix, respectively.

2.6. Cell differentiation 2.6.1. Alkaline phosphatase activity and mineralization The MSCs were seeded onto the scaffolds at a density of 50,000 cells/cm2 in osteogenic medium, consisting of growth medium supplemented with 50 mM ascorbic acid, 10 mM b-glycerophosphate, and 0.1 mM dexamethasone. After 7 or 14 days, the adherent cells on the scaffolds were lysed with Cell Lytic™ (Sigma, Germany) for 1 h, and the alkaline phosphatase (ALP) activity was quantified using a colorimetric assay kit (Abcam, UK) according to our previous report [43]. Furthermore, the total protein content of the lysate was measured using the aforementioned BCA assay. All the ALP data were normalized to the total protein content. For alizarin red staining (ARS), scaffolds were fixed in 4% (w/v) formalin for 30 min after 21 and 28 days and then washed with PBS. The samples were stained in 2% (w/v) ARS solution (pH 4.2, Sigma, Germany) for 20 min, followed by three washes with PBS. To quantify the calcium deposition, samples were incubated for 60 min with 0.2% (w/v) ARS. Subsequently 10% (w/v) cetyl pyridinium was added for 60 min to extract the calcium-bound ARS. Absorbance was measured at 595 nm and corrected at 655 nm [44].

2.4. Cell proliferation assay Bone marrow was obtained from patients undergoing orthopedic surgery (University Medical Center Utrecht, Utrecht, The Netherlands) after written informed consent and approval of the local medical ethical committee. The isolation method of MSCs from the bone marrow has been previously reported [42]. Cells were expanded in growth medium consisting of Dulbecco's modified Eagle medium (DMEM), 15% (v/v) fetal bovine serum (FBS) and 1% (v/v) penicillin/streptomycin solution (all from Invitrogen, US), to a maximum passage of 4. The nanofiber matrixes were punched with 15 mm diameter punches and put into 24well plates for sterilization under UV radiation for 1 h. They were subsequently rinsed three times with sterile PBS for 5 min and immersed in growth medium overnight. The hMSCs were seeded on the scaffolds after removing the medium. Cell proliferation was quantified with an (3-[4,5-dimethylthiazol-2yl]-2,5-diphenyltetrazolium bromide, Sigma, Germany) MTT assay. Cells were seeded at 20,000 cells/cm2 on each fiber surface for 1, 7 and 14 days in a humidified incubator with 5% CO2 at 37 °C. Then, the samples were transferred to a new plate for the MTT assay. Metabolic activity was determined by adding 50 μL of a 5 mg/ml solution to 250 μl of the culture medium at the predetermined time points. After 4 h incubation at 37 °C, 300 μl of dimethyl sulfoxide was added to the supernatant, and the absorbance was measured at 570 nm (Fluoroskan Ascent FL, Thermo Fisher Scientific, Spain). The cell viability on each scaffold was assessed by live-dead assay (Life Technologies, UK) according to the manufacturer's instructions after 3 days in the growth medium. A fluorescence microscope (Leica SP8X confocal microscopy, Germany) was used to qualitatively assess live and dead cells. The livedead images (n = 3) were quantified through ImageJ software according to Eq. (4):

Live or Dead cells% =

SL or SD × 100 ST

2.6.2. RNA isolation and gene expression The MSCs were seeded onto the composite nanofibers at a density of 50,000 cells/cm2 in osteogenic medium for 7 or 14 days. To evaluate their osteogenic activity, the expression level of bone-related markers was quantified by real-time polymerase chain reaction (RT-PCR). Total ribonucleic acid (RNA) was extracted from each type of matrix with Trizol reagent (Sigma, Germany), and the RNA concentration was measured by a NanoDrop spectrophotometer (Thermo Fischer Scientific, US). The RNA was reverse-transcribed into complementary deoxyribonucleic acid (cDNA) using the iScript cDNA Synthesis kit (BioRad, US) using the manufacturer's instructions. Finally, RT-PCR assays were conducted using the SYBR Green system (Bio-Rad, US) on an RTPCR machine (Applied Biosystems, Life Technologies, US). The PCR amplification was done under the following conditions: 95 °C for 10 min, 40 cycles at 95 °C for 10 s, and 60 °C for 60 s. For the relative quantification of target genes, the glyceraldehyde 3-phosphate dehydrogenase (GAPDH) levels were used as a reference and the 2−ΔΔCt method was applied to calculate the expression level of target genes. The expression data of the target genes were normalized by GAPDH and reported as the fold change relative to the pure PCL control group. A sequence of primers (Table 1) was used to quantify mRNA levels of GAPDH, alkaline phosphatase (ALP), runt-related transcription factor 2 (RUNX-2), osteocalcin (OCN), osteopontin (OPN), collagen type 1 and bone morphogenetic protein 2 (BMP-2).

(4)

where SL and SDis the surface area of live cells (green) and dead cells (red), respectively. sT is the surface area of total number of cells (red and green).

2.5. Cellular morphology Cell morphology was evaluated by SEM after 3 days of culturing cells in the growth medium. The MSCs were washed with PBS and fixed with 2.5% glutaraldehyde (Invitrogen, US) at room temperature for 60 min. Then, the samples were gradually dehydrated through a 50–100% ethanol series. The dry specimens were visualized with SEM as aforementioned. To visualize the actin organization, the cells were cultured on the scaffolds in growth medium for 3 days. The scaffolds were then fixed with 4% (w/v) formalin and permeabilized with 0.2% (v/v) Triton X100 (Sigma, Germany) in PBS. Subsequently, the cells were stained with 2.5 μg/ml tetramethylrhodamine B isothiocyanate (TRITC)-labelled phalloidin (Sigma, Germany) to reveal the cytoskeletal organization of actin filaments, together with and DAPI (2 μg/ml, ab104139, Abcam, UK) to stain nuclei. The samples were imaged by confocal microscopy (Leica SP8X, Germany). The cell length, cell width and cell area were measured using ImageJ software and used to quantify the mean aspect ratio (i.e., the ratio of cell length to cell width) for a total of 20 cells (4 (cells/image) × 5 images).

Table 1 Primer sequences for quantitative RT-PCR analysis. Target gene

Primer Sequence

Annealing temperature

Osteopontin (SPP1) COL1a1

F: GCCGAGGTGATAGTGTGGTT R: TGAGGTGATGTCCTCGTCTG F: ATGCCTGGTGAACGTGGT R: AGGAGAGCCATCAGCACCT F: GTGCAGAGTCCAGCAAAGGT R: TCAGCCAACTCGTCACAGTC F:GGCTGGAGATGGACAAGTTC R: CAGATTTCCCAGCGTCCTTG F: CTCATTTCCTGGTATGACAACGA R: CTTCCTCTTGTGCTCTTGCT F: ATGACACTGCCACCTCTGA R: ATGAAATGCTTGGGAACTGC F: GGAGGCAAAGAAAAGGAACGGA R: GAAGCAGCAACGCTAGAAGACA

60 °C

Osteocalcin ALP GAPDH RUNX2 BMP2

3

60 °C 60 °C 60 °C 60 °C 60 °C 60 °C

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a)

SEM Images Average= 825.16

b)

TEM Images

c)

AFM Images

316.08 nm

PCL

Frequency

25 20 15 10 5 0

100

600

1100 1600

Fiber diameter (nm)

200 nm

10 μm

F-MW CNT 0.5

Average= 760.025

264.45 nm

Frequency

25 20 15 10 5 0

200

700

1200

Fiber diameter (nm)

200 nm

10 μm Average= 601.25

259.13 nm

Frequency

F-MW CNT 1

40 30 20 10 0

100

600

1100

Fiber diameter (nm)

10 μm

100 nm Average= 636.34

309.85 nm

Frequency

F-MW CNT 2

50 40 30 20 10 0

100

600

1100

1600

Fiber diameter (nm)

10 μm

100 nm

F-MW CNT 3

Average= 794.25

318.048 nm

Frequency

30 20 10 0

300

800

1300

1800

Fiber diameter (nm)

10 μm

100 nm

Fig. 1. The topographical characterization of nanofibers with various F-MWCNT concentrations. (a) SEM images show the fiber morphology, (b) TEM images show the F-MWCNT distribution in the nanofibers and (c) AFM images illustrate the surface nanoroughness. The red arrows show the orientation of F-MWCNTs and the red circles show the F-MWCNT bundles in the nanofibers. The black arrows indicate the boundary of the nanofibers. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article.)

(Cambridge, UK). The scaffolds were fixed in 4% (w/v) paraformaldehyde in PBS for 20 min and washed twice with 0.05% (v/v) Tween/PBS for 5 min. Subsequently, the cells were blocked with 2% (w/v) BSA/PBS (Sigma, Germany) and permeabilized with 0.05% (v/v) Triton X-100

2.6.3. Immunofluorescence staining All procedures were performed at room temperature unless stated otherwise. Antibodies against human osteopontin (OPN, ab8448) and human osteocalcin (OCN, ab13418) were purchased from Abcam 4

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b)

a)

100 C-H 2943

C=O 1723

80

Porosity (%)

Absorbance (%)

F-MWCNT 3 F-MWCNT 2 F-MWCNT 1 F-MWCNT 0.5

60 40 20

PCL

0

3500

2500

1500

0

500

-1

Wavenumber (cm )

c)

*

d)

*

1

2

3

50

*** 40

1000

Ra (nm)

Fiber diameter (nm)

1500

0.5

Percentage of F-MWCNTs

30 20

500 10 0

0 0

0.5

1

2

0

3

0.5

1

2

3

Percentage of F-MWCNTs

Percentage of F-MWCNTs

f)

e) 1.0

Surface density (m2)

Protein adsorption (mg BSA/mg NF)

1000

*

0.5

**

800 600 400 200

0.0

0

0

0.5

1

2

3

0

Percentage of F-MWCNTs

0.5

1

2

3

Percentage of F-MWCNTs

Fig. 2. Nanofiber characterization. (a) FTIR spectrum, (b) Porosity of nanofiber matrix, (c) The average nanofiber diameter (d) The average surface nanoroughness (Ra) (e) The average protein adsorption on the nanofiber surface (f) The surface density of nanofibers after normalization by fiber diameter. Data are presented as the mean ± SD (n = 5); (*p < 0.05, *** p < 0.001) One-way ANOVA with Tukey correction.

(Sigma, Germany) in PBS for 1 h. The antibodies against OCN (10 μg/ ml) and OPN (10 μg/ml) were diluted in the blocking buffer and applied to the scaffolds at 4 °C overnight. After washing, the samples were incubated with Alexa Fluor 488 donkey anti-mouse IgG (10 μg/ml in blocking buffer, Invitrogen, US) and Alexa Fluor 555 donkey anti-rabbit IgG (10 μg/ml in blocking buffer, Invitrogen, US) for 1 h, respectively. The images were observed under the fluorescence microscope after counterstaining with DAPI (2 μg/ml, ab104139, Abcam, UK).

was performed as a statistical model. Values of ***p < 0.001, **p < 0.01 and *p < 0.05 were used as levels of significance. 3. Results 3.1. Characterization SEM analyses showed the formation of uniform nanofibers without the occurrence of the bead as well as a wide fiber diameter distribution (Fig. 1a). TEM measurements showed that F-MWCNTs at 0.5 and 1 wt% concentrations were completely de-bundled and aligned with the nanofiber axis (red arrows in Fig. 1b). Enhance of the F-MWCNT concentration to > 1 wt% caused bundled F-MWCNTs (red circles in

2.7. Statistical analysis All quantitative data sets are expressed as mean ± SD. One-way analysis of variance (ANOVA) with Tukey or LSD post-hoc correction 5

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***

Single nanofiber

40 30 20 10

F-MWCNT 3

F-MWCNT 2

0 PCL

F-MWCNT 3

F-MWCNT 2

F-MWCNT 1

PCL

F-MWCNT 0.5

F-MWCNT 3

F-MWCNT 2

F-MWCNT 1

0

PCL

200

***

F-MWCNT 1

400

***

50

F-MWCNT 0.5

Young modulus of nanofiber matrix (Mpa)

***

F-MWCNT 0.5

Young modulus of single nanofiber (Mpa)

600

Nanofiber matrix

Fig. 3. Mechanical characterization of nanofibers with various F-MWCNT concentrations. Young modulus of single nanofiber and nanofiber matrix. Data are presented as the mean ± SD (n = 5); (*** p < 0.001).

was larger on the higher stiffness nanofibers. The MSCs cultured on 0.5 and 1 wt% nanofibers showed more and larger focal adhesion than other groups (Fig. 4b). Both quantitative (Fig. 4c and e) and qualitative (Fig. 4d) assays showed similar cell viability in the different groups. All nanofiber structures supported the normal proliferation of MSCs up to 14 days of culture, as seen by their significantly enhanced metabolic activities (Fig. 4b). These data collectively show that the incorporation of FMWCNTs in the PCL nanofibers does not hamper MSC adherence and proliferation.

Fig. 1b) that created nanoroughness on the surface of nanofibers. The AFM results further verified the results of the TEM findings (Fig. 1c). Both TEM and AFM images confirmed that there was no nanoroughness on the nanofiber surface at 0.5 and 1 wt% concentrations, but it was only induced by 2 and 3 wt% F-MWCNT concentrations (black arrows in Fig. 1b). The FTIR spectrum showed that the F-MWCNTs were completely encapsulated in the PCL nanofibers and no additional peaks were observed in composite groups (Fig. 2a). No significant differences were observed in the fiber porosity following incorporation of F-MWCNTs in PCL nanofibers (Fig. 2b). The incorporation of F-MWCNT led to a decrease in the average diameters of the composite nanofibers up to 2 wt% in comparison with pure PCL (Fig. 2c). The average surface nanoroughness (Ra) was elevated by increasing the F-MWCNT concentration to > 2 wt% (Fig. 2d). In order to validate the AFM results, a protein adsorption assay was used. The amount of adsorbed proteins on nanofibers at a concentration above 1 wt% was higher than other groups (Fig. 2e). After normalizing the adsorbed proteins against nanofiber diameter (Eq. 2), the result showed a higher surface density at 2 and 3 wt% F-MWCNTs/PCL nanofibers which are attributed to the surface nanoroughness (compare Fig. 2d and f).

3.3.2. ALP activity and matrix mineralization ALP activity was measured in MSCs as an early marker of osteogenic differentiation. On day 3, no differences in MSC ALP activity were found between groups. On day 7 and 14, however, increased ALP activity was measured in MSCs cultured on composite nanofibers with relatively higher stiffness or nanoroughness as compared to pure PCL nanofibers (Fig. 5a). The stimulatory effect on ALP activity was most pronounced for composite nanofibers associated with the highest stiffness. ARS of composite nanofibers was performed to specifically confirm calcium deposition after 21 and 28 days. Mineralized matrix formation was observed in all nanofiber groups. However, the scaffolds with higher stiffness showed higher calcium deposition in day 21 compared to other scaffolds. After 28 days, both higher roughness and stiffness groups showed more mineralization as compared to the nanofibers with the only PCL (Fig. 5b and c).

3.2. Mechanical characterization The single nanofibers and the nanofiber matrixes showed considerable differences with respect to their stiffness (Fig. 3). The Young modulus of single nanofibers increased when adding up to 1% of the FMWCNTs. Thus, PCL with 0.5% and 1% F-MWCNTs exhibited the maximum single nanofiber modulus. Subsequently, a further increase of the F-MWCNT concentration (2 and 3 wt%) led to a decrease in the single nanofiber modulus. In contrast, there was a direct relationship between the young modulus of the nanofiber matrix and the F-MWCNT concentration.

3.3.3. Osteogenic gene expression and immunofluorescence staining The influence of nanofiber design was studied on the levels of key osteoblast gene markers in MSCs to elucidate whether different osteogenic pathways could be involved in material-induced osteogenic differentiation (Fig. 6a). In line with the measured ALP activity, the day 7 ALP mRNA levels were remarkably upregulated on composite nanofibers when compared to nanofibers with the only PCL. The highest ALP expression level was found after culture on composite nanofibers with higher stiffness, which showed a significant difference at 14 days. The expression level of RUNX-2 was higher on the scaffolds with higher stiffness and roughness compared to scaffolds, with only PCL after 14 days. In addition, there were no significant differences in the expression levels of COL1 and OCN between all groups at day 7. However, a sharp increase (9 fold) in the expression level of OCN was observed on the matrix with higher stiffness at day 14, and this expression was dramatically down-regulated on scaffolds with higher roughness. In contrast, the expression level of BMP-2 on day 7 and OPN at both time points were increased on the composite nanofibers with higher roughness. It was also observed that stiffer substrates increased the expression

3.3. Human MSC cytocompatibility and osteogenic differentiation 3.3.1. Cell cytocompatibility SEM images demonstrated comparable cell attachment and homogeneous cell distribution in the different composite nanofiber groups after three days (Fig. 4a). Cytoskeleton staining showed that the nanofiber stiffness and nanoroughness induced different cell morphologies (Fig. 4a). Analysis of individual cells revealed that the cells cultured on the higher roughness nanofibers were more elongated as compared to PCL group. Additionally, the aspect ratio was highest for the 2 and 3 wt% F-MWCNT nanofibers. In comparison, the cellular area 6

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a)

PCL

F-MWCNT 0.5

F-MWCNT 2

F-MWCNT 3

Cytoskeleton staining

SEM

F-MWCNT 1

DAPI

Phalloidin

20 μm

DAPI

Phalloidin

20 μm

DAPI

DAPI

20 μm

Phalloidin

c)

20 μm

Phalloidin

***

1.5

OD (570nm)

b)

DAPI

20 μm

Phalloidin

***

1.0

*** 0.5

Day 7

F-MWCNT 3

F-MWCNT 1

F-MWCNT 2

PCL

F-MWCNT 0.5

F-MWCNT 3

F-MWCNT 2

F-MWCNT 1

PCL

Day 1

F-MWCNT 0.5

F-MWCNT 3

F-MWCNT 2

F-MWCNT 1

PCL

F-MWCNT 0.5

0.0

Day 14

e)

d) TCP

PCL

F-MWCNT 0.5

250 μm

F-MWCNT 1

250 μm

250 μm

F-MWCNT 2

F-MWCNT 3

Cell area (%)

150

live cells Dead cells

100

50

TC P FM W PC L C F- N T M 0 W .5 F- CN T M W 1 F- CN T M W 2 C N T 3

0

250 μm

250 μm

250 μm

Fig. 4. The effect of stiffness and nanoroughness on the adhesion and proliferation of human MSCs. (a) Representative images of the SEM and cytoskeleton staining of MSCs on nanofibers with distinct stiffness and roughness after 3 days (b) Quantification of the cell aspect ratio and cell area of cytoskeleton images (n = 5) (c) MTT assay was performed after 1, 7 and 14 days of culture to show the cell proliferation (n = 3) (d) Representative images of Live-dead assay after 3 day (e) Quantification of live-dead assay. Data are presented as the mean ± SD; (*** p < 0.001).

7

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b)

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Fig. 5. ALP activity and matrix mineralization of human MSCs of composite nanofibers with distinct stiffness and roughness (a) ALP activity after 3, 7 and 14 days of cell culture in osteogenic medium (b) quantification of calcium deposition after 14, 21 and 28 days of cell culture in osteogenic medium (c) representative ARS staining of composite nanofibers after 21 and 28 days in osteogenic medium. Data are presented as the mean ± SD (n = 3); (*p < 0.05, *** p < 0.001).

MWCNT concentrations in the scaffolds. The mechanical and topographical analysis showed that the maximum single nanofiber stiffness was related to composite nanofibers with an F-MWCNT concentration of 0.5 and 1 wt%, i.e., no nanoroughness on the surface. At this concentration, the uniform and aligned dispersion of F-MWCNTs in the PCL matrix most likely helped to transfer the mechanical stress from the PCL matrix to the F-MWCNTs, which have excellent intrinsic mechanical properties [53]. On the other hand, the maximum nanoroughness was associated with an F-MWCNTs concentration of 2 and 3 wt%, with no significant differences of stiffness with only PCL nanofibers (Figs. 2 and 3). As an explanation for this observation, it is thought that an increase in the concentration of F-MWCNTs increases their tendency to aggregate into bundles due to their high intertube van der Waals forces [54]. Furthermore, the interfacial area between F-MWCNTs and the PCL matrix were dramatically reduced, which ultimately led to increase in the nanoroughness and a decrease in the stiffness [55,56]. Nanofiber roughness and diameter are equally crucial nanoscale parameters that affect cell behavior [57,58], and were successfully modulated by the current F-MWCNT incorporation method. A protein adsorption assay was used to explore the effect of these two parameters simultaneously, considering that the protein adsorption onto the material surface is highly related to the specific surface area of the matrix (i.e., nanofiber diameter and surface nanoroughness) [59–61]. After normalizing the protein adsorption results against the nanofiber diameter (Fig. 2f), it was verified that the nanoroughness is the only

level of OCN and COL1 at day 14 as well as a substrate with a higher roughness upregulated the expression level of BMP-2 and OPN. To confirm the different mRNA profiles found for the late osteogenic markers OCN and OPN, immunostaining was performed to characterize their expression at the protein level. Accordingly, OCN expression was found to be strongly related to the scaffold stiffness, while the OPN expression was clearly associated with the scaffold roughness (Fig. 6b). 4. Discussion In this study, a complete comparison was performed between the isolated effect of nanoroughness and stiffness of electrospun nanofibers on osteogenic regulation. For this purpose, carbon nanotubes were incorporated into nanofibers to fabricate scaffolds with improved physical properties. Electrospinning is a well-known technique to fabricate nanofibers for a wide range of applications [45,46]. Using carbon nanomaterial in nanofibers as a filler is a well-known method to modify the physical properties [47–49]. When using this method, the homogeneous dispersion of CNTs in small bundle size is a very critical step to improve the interfacial area between the CNTs and the polymer [50]. Based on the available literature, it was presumed that functionalization, sonication, and CNT concentration could all improve the physical properties of the nanofibers by determining the interfacial area between the CNTs and the polymer [51,52]. We, therefore, applied the FMWCNTs using an optimized sonication process and varied the PCL/F8

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Fig. 6. Extent of osteogenic differentiation on composite nanofibers with distinct stiffness and nanoroughness (a) Real-time PCR after 7 and 14 culture days in osteogenic medium, ALP/RUNX2 (early), COLI/BMP-2 (middle) and OCN/OPN (late) gene expressions (b) Representative images of OCN and OPN staining of MSCs cultured in osteogenic medium for 28 days. Data are presented as the mean ± SD (n = 3); (*p < 0.05, ** p < 0.01, *** p < 0.001).

In the current study, the carbon nanomaterial-modified biomaterials were found to elicit various beneficial responses in human bone-derived MSCs, i.e., the predominant cell source of new bone formation in vivo [66]. The PCL-only scaffolds demonstrated excellent MSC cytocompatibility, which was unaffected by F-MWCNT loading. Nevertheless, several key differences were found when comparing PCL-only and PCL/F-MWCNT scaffolds. First, the MSCs showed unique morphologies and actin organizations on nanofibers associated with different physical properties (Fig. 4a). Most noticeably, the MSCs displayed increased stress fibers and complete spreading on nanofibers with higher stiffness, while in contrast, cell elongation was observed for rougher surfaces. These findings are in agreement to other reports, showing that stiff substrates support cell spreading and actin cytoskeleton organization into stress fibers [67,68], whereas nanoscale matrix topography affects the microscale cell motility through adhesions and

parameter changed by PCL/F-MWCNTs (2 and 3 wt%). The mechanical properties of both the entire nanofiber matrix and single nanofibers are crucial factors to engineer the desired scaffold. The matrix stiffness is responsible for defining the degradation kinetics of scaffolds [62], whereas single nanofibers are in charge of transferring the applied extrinsic force to cells [63]. Increasing evidence points to the fact that the tensile modulus of a single nanofiber may be harnessed for increased osteogenesis to show the effective elasticity sensed by the cells [64,65]. By performing separate studies on nanofiber matrixes and single nanofibers, we were able to demonstrate that 0.5 and 1 wt% FMWCNTs increased the single nanofiber stiffness, while, the matrix modulus increased once the F-MWCNTs increased. These differences can be due to the randomness structure of nanofibers, cohesion among nanofibers at contact points and high variation of nanofiber diameters in the matrix [26]. 9

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actin organization [12,69]. Second, we confirmed by measuring a set of bone markers that the provided physical cues can play an essential role to improve its pro-osteogenic capacity, i.e., bone regeneration. In particular, an increase in either nanoroughness or stiffness both enhanced the expression of late osteogenic markers, with a higher stiffness promoting osteogenesis relatively earlier than a rougher substrate. Based on an earlier study [12], it seems that the AKT/YAP signaling pathway played an important role in stiffness-mediated osteogenesis on electrospun scaffolds, as was seen here. In contrast, a different mechanism seems to induce topography-mediated mechanotransduction, and which requires the formation of focal adhesion points [18,70,71]. Irrespective of the underlying mechanism, the MSCs were found to be more sensitive to stiffness than nanoroughness in their osteogenic regulation [11]. F-MWCNT incorporation into electrospun nanofibers can, therefore, be considered a useful tool to develop new bone substitute biomaterials without changing the surface chemistry. Remarkably, a higher substrate nanoroughness or stiffness independently improved the expression level of the osteogenic markers OPN and OCN, respectively. The different role of these proteins in osteogenesis leads to the suggestion that material stiffness and topography activate unique signaling pathways leading to osteogenesis, which needs more investigation in future studies. As compared to OPN, OCN is expressed later during osteogenic differentiation and has a more indispensable role in the mineralization stage [72–74]. Furthermore, considering the various roles of OPN in the regulation of macrophage activity and osteoclast formation [75,76], two processes highly important for bone formation, MSC-macrophage coculture can potentially reveal certain osteo-immunomodulatory effects following F-MWCNTmediated mechanotransduction.

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5. Conclusions The incorporation of F-MWCNTs into PCL nanofibers was found to be an effective method to independently control the material surface nanoroughness and stiffness, two crucial parameters involved in regulating cell function. The composites showed excellent cytocompatibility. Moreover, as a result of improved surface nanoroughness or stiffness, enhanced osteoblast differentiation was measured for a wide panel of markers. Changes in nanoroughness and stiffness affected the matrix mineralization and the expression of the late osteogenic markers differently, suggesting that these physical cues regulate osteogenesis through different mechanisms. In conclusion, the current method allowed managing of the single nanofiber stiffness and nanoroughness to modulate and promote the osteogenic response. This can help our further understanding of the mechano-biological responses in osteoprogenitor cells, and could ultimately lead to improved design of bone tissue engineering scaffolds. Declaration of competing interest There are no conflicts to declare. Acknowledgment The supports of Dr. George Posthuma from Cell Microscopy Core of UMC Utrecht in TEM analysis is acknowledged. References [1] M.P. Lutolf, F.E. Weber, H.G. Schmoekel, J.C. Schense, T. Kohler, R. Müller, J.A. Hubbell, Repair of bone defects using synthetic mimetics of collagenous extracellular matrices, Nat. Biotechnol. 21 (5) (2003) 513–518. [2] M.A. Fernandez-Yague, S.A. Abbah, L. McNamara, D.I. Zeugolis, A. Pandit, M.J. Biggs, Biomimetic approaches in bone tissue engineering: integrating biological and physicomechanical strategies, Adv. Drug Deliv. Rev. 84 (2015) 1–29. [3] D. Bhattarai, L. Aguilar, C. Park, C. Kim, A review on properties of natural and synthetic based electrospun fibrous materials for bone tissue engineering,

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