Enhanced Osteogenic Differentiation of Stem Cells via Microfluidics Synthesized Nanoparticles Mohammad Mahdi Hasani-Sadrabadi, Sana Pour Hajrezaei, Shahriar Hojjati Emami, Ghasem Bahlakeh, Leila Daneshmandi, Erfan Dashtimoghadam, Ehsan Seyedjafari, Karl I. Jacob, Lobat Tayebi PII: DOI: Reference:

S1549-9634(15)00094-5 doi: 10.1016/j.nano.2015.04.005 NANO 1114

To appear in:

Nanomedicine: Nanotechnology, Biology, and Medicine

Received date: Revised date: Accepted date:

28 January 2015 25 March 2015 8 April 2015

Please cite this article as: Hasani-Sadrabadi Mohammad Mahdi, Hajrezaei Sana Pour, Emami Shahriar Hojjati, Bahlakeh Ghasem, Daneshmandi Leila, Dashtimoghadam Erfan, Seyedjafari Ehsan, Jacob Karl I., Tayebi Lobat, Enhanced Osteogenic Differentiation of Stem Cells via Microfluidics Synthesized Nanoparticles, Nanomedicine: Nanotechnology, Biology, and Medicine (2015), doi: 10.1016/j.nano.2015.04.005

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ACCEPTED MANUSCRIPT Enhanced Osteogenic Differentiation of Stem Cells via Microfluidics Synthesized Nanoparticles Mohammad Mahdi Hasani-Sadrabadia,b,c, Sana Pour Hajrezaeib, Shahriar Hojjati Emamib,

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Ghasem Bahlakehb, Leila Daneshmandid, Erfan Dashtimoghadame, Ehsan Seyedjafarif, Karl I. Jacoba,*, Lobat Tayebie,g,*

Parker H. Petit Institute for Bioengineering and Bioscience, G.W. Woodruff School of

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a

Mechanical Engineering and School of Materials Science and Engineering, Georgia b

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Institute of Technology, Atlanta, 30332-0295 GA, USA.

Center of Excellence in Biomaterials, Department of Biomedical Engineering and

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Department of Chemical Engineering, Amirkabir University of Technology, Tehran, Iran. Department of Stem Cells and Developmental Biology at Cell Science Research Center,

Royan Institute for Stem Cell Biology and Technology, ACECR, 19395-4644, Tehran,

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Iran.

Department of Bioengineering, Temple University, Philadelphia, 19122 PA, USA.

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Department of Developmental Sciences, Marquette University School of Dentistry,

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d

Milwaukee, 53201 WI, USA.

Department of Biotechnology, College of Science, University of Tehran, Tehran, Iran.

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Biomaterials and Advanced Drug Delivery Laboratory, Stanford University School of

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Medicine, Palo Alto, 94304 CA, USA.

*Corresponding authors: (K.I.J.) [email protected]; (L.T.) [email protected]. Word count: Abstract: 150 words Manuscript word count: 4839 words in the body text + 650 words in the seven Figure legends. Number of references: 35 Number of Figures: 7 Figures

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ACCEPTED MANUSCRIPT Abstract Advancement of bone tissue engineering as an alternative for bone regeneration has

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attracted significant interest due to its potential in reducing the costs and surgical trauma

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affiliated with the effective treatment of bone defects. We have improved the conventional approach of producing polymeric nanoparticles(NPs), as one of the most

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promising choices for drug delivery systems, using a microfluidics platform, thus further

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improving our control over osteogenic differentiation of mesenchymal stem cells. Molecular dynamics simulations were carried out for a theoretical understanding of our

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experiments in order to get a more detailed molecular-scale insight into the drug-carrier interactions. In this work, with the sustained intracellular delivery of dexamethasone

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from microfluidics-synthesized NPs, we explored the effects of particle design on

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controlling stem cell fates. We believe that the insights learned from this work will lead

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cell therapeutics.

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to the discovery of new strategies to tune differentiation for in situ differentiation or stem

Keywords: Polymeric nanoparticles; Microfluidics platform; Drug delivery; Osteogenic differentiation;

Background Although auto- or xenogeneic bone grafts are currently one of the most effective clinical methods for treating bone defects, but they have severe shortcomings such as pain and cosmetic issues at the donor site as well as time-consuming nature of this process due to the additional requirement of surgery.1, 2 Stem cell-mediated therapy has been suggested 2

ACCEPTED MANUSCRIPT as a new strategy for bone regeneration.3 Mesenchymal stem cells’ (MSCs) self-renewal capacity along with their ability to differentiate and produce progenitors of several cell

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types when exposed to the right factors helps in developing osteoblasts.4 Controlled

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delivery of Glucocorticosteroids (GCs) is known to regulate bone formation based on in vitro experiments. Dexamethasone (Dex) is one of the GCs that modulates osteoblast

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differentiation and mineralization through its effects on the expression and function of

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epithelial sodium channels (ENaCs). GCs bind and activate a cytoplasmatic glucocorticoid receptor. A variety of carriers have been investigated to overcome the

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adverse side-effects of free Dex delivery and also to study their internalization and ability to differentiate MSCs into osteoblasts.5-7 Nanoparticles (NPs) are one of the most

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attractive carriers for drug delivery systems (DDS). Until now, several polymeric NPs

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have been used as carriers for Dex. Among them only a few involve loading Dex in Chitosan (CS) nanoparticles. Consistent with our previous work, we are interested in

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using Chitosan as the biopolymer for synthesizing our nanoparticles.8-11 Cross-linking of

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Chitosan takes place by covalent cross-linkage with itself or from ionic gelation methods.12 In ionic gelation methods the most common cross-linking agent for Chitosan is sodium tripolyphosphates. However, for the first time we have proposed cross-linking Chitosan via adenosine triphosphate (ATP) to form CS-based NPs through ionic gelation. The release of ATP molecules during incubation with cells influences cellular viability positively. Current fabrication methods for nanoparticles such as conventional bulk mixing result in extremely polydisperse NPs that exhibit various physicochemical properties which are difficult to control.8, 13 Moreover these methods often require further post-processing steps.14 In the present study we have used microfluidics assisted

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ACCEPTED MANUSCRIPT fabrication as a bottom-top method to synthesize our nanoparticles.8-11,

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advantage of microfluidics is the high control over physical characteristics by simply

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adjusting the flow parameters.15 By applying small changes in flow parameters we can

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make nanoparticles of different characteristics that are applicable in various fields, making this method to be a very attractive approach for nanoparticle production.16, 17

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Although, the focus of the present paper lies in deeper in vitro studies investigating the

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potential applications of the microfluidics technique for NP synthesis we have gone a step further in reporting the osteogenic efficiency of Dex loaded from these

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microfluidics-assisted prepared NPs. To demonstrate the effects of Dex on osteogenic differentiation of MSCs, we analyzed the established markers of the osteoblastic

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phenotype, including alkaline phosphate activity, osteocalcin, and calcified matrix

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deposition.

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Methods

Microfluidics device fabrication: The microfluidics device was fabricated with poly (dimethylsiloxane) (PDMS) using a standard micromolding process as described in Supplementary Information. Characterization details of nanoparticles are mentioned in Supplementary Information. MSCs were isolated from the femurs and tibias of 30-dayold rats (see Supplementary Information (SI)). All experiments involving animals were in compliance with Provisions and General Recommendation of Experimental Animals Administration.

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ACCEPTED MANUSCRIPT To evaluate cellular internalization of our NPs, MSCs cells were cultured with FITClabeled NPs and subsequently analyzed using flow cytometry on a CyAn ADP Analyzer

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(FACS, Beckman Coulter, Inc) and ZEISS LSM700 UP2 (Jena, Germany) confocal laser-

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scanning microscope (CLSM) (see SI). The cytotoxicity tests were assessed using an MTT colorimetric assay. Cytotoxicity of the nanoparticles was determined after 72 h

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incubation with MSCs cultured in DMEM medium (Sigma-Aldrich, St. Louis, MO,

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USA) at 37°C in a 5% CO2 incubator. To determine cytotoxicity, cells were plated at a density of 10,000 per well in 96 well plates and then incubated overnight. Reverse

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transcription polymerase chain reaction (RT-PCR) was used to analysis the mRNA levels of alkaline phosphatase (ALP), core binding factor alpha l (cbfa1) and the alpha l chain of

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collagen type I (COL1A1) on day 7 of differentiation (see SI). To quantify ALP activity,

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cell culture supernatants were collected on day 14 of differentiation, centrifuged at 2000 g for 10 min at 4°C to remove any debris and analyzed for ALP using an ALP Kit

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according to the manufacturer’s instructions. The absorbance was read at 405 nm using a

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spectrophotometer after adding the reagents. The amount of Osteocalcin secreted to the culture medium was measured using an ELISA plate reader (Safire II, Tecan Sales Switzerland AG, Mannedorf, CH). The ELISA assay was based on two goat Osteocalcin antibodies, recognizing both the carboxylated (C-terminal protein) and decarboxylated (N-terminal protein) rat osteocalcin as reported by Santos et al.18. Samples were collected at days 2, 7 and 14 of culture period and stored at -18°C for later quantification. Samples were run in triplicates based on the manufacturer’s instructions and compared against rat osteocalcin standards. On day 21 of incubation, the amount of secreted calcium was quantified using the O-Cresolphthalein Complexone method (Sigma-Aldrich, St. Louis,

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ACCEPTED MANUSCRIPT MO, USA) as mentioned before.19 The amount of calcium present was normalized to the total protein content of the cells (μg.mg-1). The presented data were expressed as average

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± SD. The results were statistically analyzed using t-tests. For all tests the statistical

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significance was set at p < 0.05. p < 0.01 and p < 0.001 are considered as very and

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extremely statistically significant, respectively.

Results

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We used microfluidics devices based on hydrodynamic flow focusing on producing monodisperse nanoparticles with high control. As shown in Figure 1a, our system is

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based on a cross-junction PDMS chip consisting of two inlets for ATP-containing water,

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one outlet for the Chitosan and Dex solution and one outlet for extracting the fabricated NPs. We used micropumps to change the flow ratio (FR) of the Chitosan and ATP

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streams from 0.03 to 0.2 and adjust the degree of flow focusing which resulted the

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mixing time (τmix) to be in the range of milliseconds. According to Supplementary Eq. (S1), by changing the FR of the Chitosan and ATP streams from 0.03 to 0.2, any mixing time between 2.5 and 75 ms can be achieved. The hydrodynamic diameter of prepared NPs are characterized using dynamic light scattering (DLS) and shown in Figure 1b. The DLS results were confirmed by TEM images displayed in Figure 1c-e showing a similar size and morphology for Dex-loaded NPs. As seen in Figure 1b, the on-chip synthesized NPs are smaller than bulk NPs, achieving smaller particle sizes at lower flow ratios. Nanoparticles prepared with the microfluidics platform also have a much narrower size distribution with the shortest τmix resulting in the most uniform particles.

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ACCEPTED MANUSCRIPT To better understand the effect of mixing time on the physical characteristics of our NPs, compactness was calculated from the local polymer concentration (CNP) inside the NPs

3ct  2 1  2cNP Rh  wcNP

  1 1  cos  wcNP    (1) sin  wc    wcNP  

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

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according to Eq. (1),20

where τ and ct stand for the turbidity and total polymer concentration of the NP w  4 Rh  dn / dc  /  no , wherein Rh is the hydrodynamic

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suspension, respectively.

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radius of particles, γ is the wavelength for turbidity measurements, no is the refractive

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index of the solvent and dn/dc is the refractive index increment of the polymer. The CNP values calculated from Eq. (1) for the synthesized NPs, composed of Chitosan

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chains and cross-linked with ATP molecules, are shown in Figure 1f for different sizes.

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The on-chip synthesized nanoparticles are more compact than those formed through bulk mixing, leading to higher turbidity levels in solutions of the same concentration. In

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addition, we observed that increasing the size of microfluidics synthesized NPs decreased compactness.

This is a unique ability of our microfluidics device to produce more compact nanoparticles especially at lower mixing times (Figure 1f). It is also worth mentioning that with the hydrodynamic flow focusing approach employed in this work, we can tune particle compactness by altering the size of our μF-NPs depending on the required application. Having the values of hydrodynamic size and CNP of the NPs, molecular weight of the spherical particles, MNP, was calculated20 according to Eq. (2),

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ACCEPTED MANUSCRIPT 4 M NP   Rh3CNP N A (2) 3

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The calculated MNP values are displayed in Figure 1c. The aggregation number of the

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polymer chains in the NPs, Nagg, are determined based on Eq. (3) and are shown in Figure

M NP (3) Mn

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N agg 

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1g.

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where Mn is the numerical average molecular weight of the polymer. As seen in Figure 2, MTT-based cell viability assay of MSCs after 72 h of exposure to

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nanoparticles at 37 °C showed the empty nanoparticles synthesized using microfluidics

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platform or through bulk mixing to be non-toxic, even at high concentrations of 800 nM. We observed cell viability to increase with decreasing the NP size and to be independent

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of NP concentration.

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To evaluate cellular internalization of our NPs, MSCs were cultured for 2 h with FITC labeled NPs. After culture, the cells were rinsed with EDTA (pH 5.0) to remove surface bound NPs, and then quantified for fluorescence intensity and therefore degree of NP uptake with FACS (Figure 3b). Results indicated that the NPs prepared with the microfluidics platform were internalized more than the bulk synthesized NPs. This is thought to be due to their higher surface charge and smaller hydrodynamic size. The fluorescence intensity of the cells cultured with microfluidics NPs (μF-NP-58 nm) was two folds higher than that of the cells cultured with bulk mixing NPs, with lower intensity levels as particle size increases (Figure 3).

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ACCEPTED MANUSCRIPT After achieving particles with controllable physical characteristics, the ability of our microfluidics platform to encapsulate an anti-inflammatory agent, Dexamethasone (Dex),

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was evaluated. As stated above, our aim is to investigate the ability of these Dex-loaded

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CS-NPs to induce MSC differentiation. Dex-loaded NPs are consistently larger than their unloaded counterparts. With the addition of Dex, however the NP size was still

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maintained in the acceptable size range for passive targeted drug delivery, allowing these

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NPs to serve as potential agents for in situ differentiation. The in vitro release profile of Dex-loaded NPs is presented in Figure 4. Results are based on Dex release in PBS

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(phosphate buffer saline) at pH of 7.4 over three weeks (see Supplementary Information for details). The diffusion coefficient of Dex molecules within the NPs, was also

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investigated through the fraction of released drug and the square root of time according to

1/2

(4)

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Mt  D   6  M R 

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Eq. 4,

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Where Mt/M∞ represents the fraction of released drug at time t, D is the diffusion coefficient of the drug molecules and R is the radius of the nanoparticles (based on DLS results). The required times to release half of the loaded drugs (t50) from the nanoparticles is shown in Figure 4b. Molecular dynamics (MD) simulations were conducted on our experiments to get a detailed molecular-scale insight into the interactions of Dex with the CS nanoparticles. To perform MD simulations on Dex encapsulated CS-NPs, a 3D amorphous cell consisting of Chitosan chains encapsulating the Dex drug molecules was constructed. Ten chains of Chitosan with a polymerization degree of 10, and four (quantum mechanics9

ACCEPTED MANUSCRIPT optimized) Dex molecules were used to generate the 3D simulation cell. The final 3D simulation cell was subjected to MD simulations, which were all executed using

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GROMACS simulation code (see the Supplementary Information). The final equilibrated configuration of Dex drug molecules encapsulated by CS

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polymeric chains resulting from the last step of MD simulations is displayed in Figure 5a. Compared to the initial structure of Dex where the drug molecules were in vicinity of

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each other, in the final snapshot of the CS-Dex system they had moved towards the CS

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chains, indicating their affinity towards the CS chains in the drug-loaded nanoparticle. In order to provide a quantitative analysis of the interaction and binding affinity of Dex with

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its surrounding CS chains, non-bonded CS-Dex interaction energies were evaluated during the last 10 ns of MD simulations (Figure 5b). It was noted that all electrostatic,

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van der Waals (vdW) and interaction energies were negative, further confirming the

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strong tendency of Dex drugs to interact with their surrounding CS chains. It was also observed that vdW interactions are more than twice of that of electrostatic ones in such

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systems suggesting vdW interactions to have a significant role in the overall behavior of CS-Dex interactions. These strong interactions existing between Dex drug molecules and CS polymeric chains could lead to loading higher amounts of Dex in the CS nanoparticles, as well as lower rates of drug release, as is evident in the Dex diffusion coefficients in Dex-loaded CS-NPs for our experimental data. In order to achieve a deeper understanding on the binding behavior of Dex to CS chains in the molecular-scale, further structural evaluations were carried out considering the interactions of the functional groups of Chitosan with Dex drugs. For this purpose, radial distribution function (RDF) analyses were performed for the various functional groups of 10

ACCEPTED MANUSCRIPT CS which contribute to the occurring CS interactions with Dex. RDF  g A B  r   indicates the probability distribution of finding B atoms at distance r from the given A atoms as

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reference, and is calculated21 according to Eq. (5),  nB   N B  g A B  r       (5) 2  4 r r   V 

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Where nB stands for the number of B atoms around A atoms located inside a spherical

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shell with the thickness of ∆r, NB is the total number of B atoms during MD simulations, and V is the cell volume.

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RDF examinations were carried out for -NH2 amine, -OH hydroxyl attached to the backbone (OH1), and -OH hydroxyl in the hydroxymethyl side group of CS chains

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against the double bonded oxygen (O2) and the -OH hydroxyl group in Dex structures, as

presented in Figure 5c.

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displayed in supplementary Scheme S1. The RDF results from MD simulations are also

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All calculated RDFs demonstrate a high peak at first, which implies strong correlations

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between the CS functional groups and Dex molecules. Furthermore, it can be seen that in all RDFs the first peak’s position occurs at a distance less than 3.5 Å, which reveals the ability of CS functional groups to have hydrogen bonding (H-bonding) interactions with Dex drugs. Figure 5d presents the evolution of the average number of total H-bonds between Dex and CS chains over time. When assessing H-bonds, two geometrical criteria were considered: first the donor-acceptor distance should be less than 3.5 Å, and second the donor-hydrogen-acceptor angle should be more than 150°. To obtain a detailed evaluation, the contribution of the H-bonds of –NH2, -OH1, and –OH2 in CS encapsulating Dex was characterized (presented in Figure 5e). Results clearly

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ACCEPTED MANUSCRIPT demonstrated that a major contributor to the H-bonding interactions between CS and Dex (Figure 5d) was the H-bonding interactions of –NH2 with Dex molecules, which agrees

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well with its more pronounced peak height compared to the other RDF data in Figure 5c.

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These characteristics reveal the significant role of CS amine groups and their interactions with drug molecules such as Dex. Hence, MD simulations for the RDF data and H-bonds

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formed between CS and Dex confirm the strong interactions found in CS-Dex systems as

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well as justifying the prolonged release behavior from CS-based NPs. The effect of NP size on osteoblast-related gene expression of MSCs is presented (Figure

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6). As shown in Figure 6a, Dex treatment enhanced expression of ALP, cbfa 1 and COL1A1 genes which were absent in untreated cells. The increment level for

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microfluidics-assisted prepared NPs was significantly higher compared to bulk

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synthesized NPs.

To further investigate the effect of Dex on osteoblast differentiation, alkaline phosphate

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(ALP) activity was measured as an early osteogenic marker. ALP activity per protein

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content of MSCs cultured in different culture media for time points of 2, 7 and 14 days, indicate that Dex-loaded μF-NPs have a highly increased ALP activity compared to free Dex μF-NPs (Figure 7a). The average ALP activity per protein content in cells cultured with Dex-loaded μF-NPs was significantly higher than that for control samples (control + Dex). Osteocalcin (OC) content is a late osteogenic and bone-specific marker, which is produced right before and during matrix mineralization. Figure 7c represents the amount of osteocalcin normalized by the DNA content of MSCs cultured in different culture media containing different sizes of Dex-loaded μF-NPs for periods of 2, 7 and 14 days.

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ACCEPTED MANUSCRIPT These results are consistent with the previous ALP assay where Dex-loaded microfluidics-assisted formed CS-NPs demonstrated a strong ability to induce

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differentiation of MSCs into osteoblasts. It is worth mentioning that although no

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significant distinction was observed in the osteocalcin content of MSCs supplemented

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with different NP sizes, samples at day 7 and 14 displayed considerable difference. Calcium phosphate (CaPs) is known as a major constituent of bone tissue. To further

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examine osteogenic differentiation of MSCs in exposure to the Dex-loaded microfluidics

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produced NPs, the calcium content on the extracellular matrix was assessed (Figure 7d). The calcium deposition assay followed a similar trend to the ALP and OC assays,

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showing MSCs exposed to Dex-loaded μF-NPs to accumulate significantly higher levels

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of calcium in comparison to the control (with-/out Dex) groups. Discussion

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Nanoparticle formation here is based on ionic gelation of Chitosan and inter- and

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intermolecular cross-linking between the positive groups of Chitosan (protonated amine groups) and the phosphate groups of ATP molecules. Based on DLS results the hydrodynamic diameter of nanoparticles can be tuned by changing the mixing rate as was observed that increasing τmix led to larger particles. The size of our microfluidics-assisted fabricated nanoparticles is among the lowest reported values for medium molecular Chitosan particles. Aggregation time (gelation) (τagg) for the Chitosan-ATP systems is estimated to be around 20-30 ms. When τmix < τagg, NP formation is kinetically restricted, preventing particle growth, and ultimately producing smaller and more homogeneous and monodisperse nanoparticles than those prepared through bulk mixing techniques.11

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ACCEPTED MANUSCRIPT Moreover, the polydispersity index (PDI) of our NPs synthesized through microfluidics is 0.4. This

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high control over particle formation is attributed to the laminar mixing regime in

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microfluidics devices, unattainable in bulk mixing methods where turbulent mixing results in a broad distribution of residence times. Furthermore, our microfluidics method

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of producing Dex-loaded Chitosan nanoparticles is reproducible and independent of user

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skill.

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Based on Figure 1a, flow ratio of about 0.12 is considered as the transition point between the rapid (τmix < τagg) and slow (τmix > τagg) mixing regimes. In rapid mixing regimes the

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NPs are kinetically locked22 during formation which leads to smaller NPs that are more monodisperse and compact than those produced through bulk mixing. Based on DLS data

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increasing in τmix resulted in larger particles. According to zeta potential measurements,

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the surface charge of NPs formed using microfluidics platforms is higher than those synthesized with bulk mixing with surface charge increasing as flow ratio increases.

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Since the optimum surface charge for enhancing NP endocytosis is reported to be between 10-15 mV, our microfluidics NPs are considered suitable for intracellular delivery as their surface charge lies within this range. Our aim was to use these NPs to encapsulate Dex and release it in a sustained profile in order to promote osteogenic differentiation of MSCs. Since the cell receptor for Dex is located inside cells, these microfluidics NPs are a promising carrier for intracellular delivery of Dex molecules, to ultimately to affect cellular fate. The aggregation number of polymer chains implies aggregation of the Chitosan chains into NPs which is generally governed by local ionic molecular interactions. Here, in 14

ACCEPTED MANUSCRIPT microfluidics mixing, NP formation results from rapid aggregation through ionic gelation of Chitosan chains with ATP molecules. As seen in Figure 1g, Nagg for μF-NPs is lower

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than nanoparticles formed through standard bulk mixing. These results indicate that our

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method of producing NPs, at lower flow ratios, i.e., τmix < τagg, leads to a lower number of

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polymeric chains inside each nanoparticle.20

ATP molecules are a well-known energy source for cells, involved in almost all the

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energy-required processes inside cells, either directly or indirectly. Tissue loss may cause

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ischemia, hypoxia, reduced blood flow or disturbed oxygen supply, which can all be improved with the delivery of ATP towards the damaged tissue area. However, due to the

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short half-life of free ATP molecules in blood and its anionic nature, the direct use of ATP as a bioenergetics source has been restricted.23 Up until now the efficiency of

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various carriers and especially liposomes have been investigated for ATP delivery.

culture media.24,

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Results have indicated that administration of ATP increases the viability of cells in 25

Here the ATP molecules are used as crosslinking agents for the

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nanoparticles, aiming to increase their circulation time and thus enhance their overall efficiency as confirmed by pharmacokinetic (PK) results (see Figure SI-3 and Table SI1). As is shown in Figure 2, the enhanced cellular viability is due to cellular uptake and intracellular release of ATP molecules due to the degradation of CS-NPs at low lysosomal pH. A further advantage of our µF-NPs is their smaller size, which increases their intracellular trafficking. In bone tissue engineering applications where blood supply remains a challenge, ATP delivery through this technique can potentially provide the energy source needed to compensate for the shortage of nutrients and oxygen, keeping the cells alive and accelerating bone regeneration.

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ACCEPTED MANUSCRIPT Biomaterials alone are unable to resemble the extracellular matrix (ECM) as they lack the dynamic nature of tissues. For this reason, several reports have suggested adding

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microenvironmental factors to facilitate cellular behavior.26 Signaling biomolecules

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provide pathways to influence many of the responses involved in tissue regeneration such as cell adhesion and proliferation, as well as tissue-specific differentiation.27 Bone is one

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of the tissues, which has widely been studied for sequential delivery of biomolecules in

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order to affect in vitro cellular responses as well as in vivo bone regeneration. Many polymeric carriers have been developed to improve the release profile of biomolecules.28

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Osteogenic differentiation and subsequent processes such as calcium deposition, mineralization and remodeling progress over weeks and months. Therefore, materials

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containing the signaling molecules that activate MSCs in a controllable manner through

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regulating the signaling cascade, could influence specific cellular responses such as proliferation and differentiation. Dex is one of the members of the glucocorticoid family

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which is anti-inflammatory and immunosuppressive,29 both critical properties required in

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wound sites where inflammation during tissue repair can delay the regeneration process.30 The initial relatively rapid drug release from NPs is a result of dissolution and diffusion of the drug molecules not loaded compactly in the cores of the NPs, whereas the delayed releases can be attributed to diffusion of the drug from inside the NPs. As is also shown in Figure 4a, the microfluidics-assisted formed NPs have a slower release rate than the bulk synthesized NPs. Within 21 days, about 80%, 74%, 86%, and 98% of Dex was released from μF-58 nm, μF-74 nm, μF-93 nm, μF-125 nm NPs, respectively, whereas the amount of Dex released from bulk synthesized NPs was 99%. Our final goal in this study is to differentiate MSCs into the osteoblast phenotype. Since the differentiation is a

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ACCEPTED MANUSCRIPT process that takes place over a long continuous time, this prolonged controlled release is advantageous. We observed the particle release profiles to be dependent upon particle

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size. Normally, decreasing NP size will result in faster drug release, however because of

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the more compact nanostructure of microfluidics NPs (Figure 1f), the rate and amount of released Dex decreased with decreasing NP size (Figure 4a), which is quite desirable for

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osteogenic differentiation of MSCs. Normally, larger NPs with the same compactness

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exhibit slower release profiles, due to lower surface to volume ratios and longer diffusion distances for drugs to leave the nanoparticle core. As mentioned before, microfluidics

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allows for high control over compactness and therefore, as observed, can tune the release rates. Moreover, all nanoparticles released their Dex content in three weeks, whereas

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smaller microfluidics-assisted synthesized nanoparticles were able to keep their cargo for

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much longer (Figure 4). Diffusion coefficient and t50 values (Figure 4b) are in agreement with the other compact nanostructures of the nanoparticles formed in the rapid mixing

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regime. Therefore, these results confirm the important role of assembly time in the

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loading and release behavior of the nanoparticles. Previously, it has been reported that exposure of mesenchymal stem cells to Dex is effective in their differentiation to osteogenic cells in vitro.4, 31, 32 In regards to evaluating the effect of epithelial sodium channels (ENaCs) specific antagonist amiloride changes on osteoblast proliferation, differentiation and mineralization, Li et al. showed Dex to have a dose-dependent effect on the expression of differentiation markers. They concluded the best concentration of Dex to have a two-fold increase on cbfal and less than twofold increase on ALP expression, compared to the control which was 1×10-8.32 Moreover, it has been reported that continued release of Dex over a prolonged time has a

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ACCEPTED MANUSCRIPT positive effect on cell viability in addition to influencing the differentiation rate of MSCs.33-35 Several groups have shown that delivery of Dex from various polymeric

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Rickard et al. investigated the beneficial effects of simultaneous

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mineralization.4,

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matrices affects the expression of the genes involved in osteoblast differentiation and

delivery of Dex and BMP on expression of MSC osteogenic markers.4 Nuttelman et al.

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developed an injectable poly(ethylene glycol) (PEG)-based hydrogel containing Dex to

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study the osteogenic differentiation of human mesenchymal stem cells (hMSCs). ALP and cbfa1 gene expression levels were studied in two groups of hMSCs cultured in the

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presence of Dex released from the hydrogel and 100 nM Dex, and then compared to hMSCs cultured in control media containing no Dex. Results demonstrated that the

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maximum ALP levels of hMSCs cultured with released Dex and 100 nM Dex were 28-

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folds and 13-folds higher than cells cultured in control media, respectively. Moreover, the maximum cbfa1 levels in hMSCs cultured in the presence of released Dex and 100 nM

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Dex were 2.9-folds and 2.3-folds higher than hMSCs cultured in control media.6 In this

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context, Koehler et al. also suggested that the controlled release of Dex from a PEG hydrogel network induces osteogenic differentiation of hMSCs. According to their results, ALP activity and mineralization were significantly higher in Dex releasing hydrogels compared to hydrogels free of Dex, with ALP activity to be six-folds higher.7 Oliveira et al. reported similar results of enhancing osteogenesis by increasing extracellular matrix ALP activity and mineralization using Dex-loaded CMCht/PAMAM dendrimer nanoparticles combined with hydroxyapatite.5 Here we have investigated the effect of prolonged Dex release from nanoparticles of different sizes on the

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ACCEPTED MANUSCRIPT differentiation of MSCs into the osteoblast linage, particularly because the process of MSC osteogenesis needs a specific period of time to occur efficiently.

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Our results are consistent with previous reports,7 demonstrating that Dex markedly increases the expression of osteogenic genes. Image analysis of osteoblast-related

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markers also revealed that cells exposed to Dex-loaded μF-NPs stimulate the expression of ALP, cbfa1 and COL1A1 compared to samples containing only Dex or bulk-NPs

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without Dex. Dex-loaded μF-NPs were also revealed to significantly increase the

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expression of these osteogenic genes compared to Dex-loaded bulk-NPs. It has been proven that the size scale of biomaterials plays a major role on the fate of stem cell

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differentiation.1 Accordingly the probability of nanoparticle endocytosis is higher as is the final amount of uptake by MSCs (shown in Figure 3). Furthermore, mRNA levels

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started to decrease as the size of nanoparticles increased. Therefore, a higher amount of

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NP-released-Dex enters the cells leading to attachment of Dex to its intracellular receptor and eventual expression of osteogenic markers at a higher level.

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Based on the provided results in Figure 7, among Dex-loaded μF-NPs, smaller NPs lead to more ALP activity. As mentioned before, among the Dex-loaded μF-NPs, the smaller NPs can internalize more easily into cells, and therefore release a higher amount of Dex inside them, leading to increased osteogenic differentiation of MSCs which in turn results in higher ALP activity (Figure 7b). Our results indicate that the Dex-loaded microfluidics-synthesized Chitosan NPs are beneficial in differentiating MSCs since they have proven to promote early osteogenic differentiation at higher levels than those cells cultured in medium without Dex-loaded NPs.

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ACCEPTED MANUSCRIPT Here we proposed a fine-tuning method for producing Chitosan nanoparticles using a PDMS microfluidics device. We were able to control the main physical properties of the

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resulted nanoparticles including size, surface charge and cargo release rate, all by

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increasing the mixing rate. To investigate the in vitro behavior of our microfluidics NPs, the effect of Dex-loaded NPs on the osteogenic differentiation of MSCs was examined.

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Furthermore, to achieve a molecular level understanding on our experimental results,

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molecular dynamics simulation techniques were applied to study the interaction of the Dex drug with the Chitosan biopolymer. Evaluations of Dex binding affinity in terms of

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interaction energy, RDF and hydrogen-bond number revealed strong interactions between Dex and Chitosan chains, which were consistent with the lower Dex diffusion

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coefficients for release of Dex molecules. We observed these microfluidics-synthesized

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NPs have a smaller size compared to conventional bulk synthesized NPs and can induce differentiation more since they allow Dex to be released inside the cells, and therefore are

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more efficient in controlling the cellular fate. In summary, we have developed a system,

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which can be used for synthesizing drug-loaded polymeric nanoparticles or more generally nanoparticles for applications beyond drug delivery and as a new tool which can be applied for cells and tissue engineering strategies.

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ACCEPTED MANUSCRIPT Figure Legends Figure 1. The physical characteristics of microfluidics-synthesized nanoparticles. (a)

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Schematic representation of our cross-junction microfluidics device used to

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hydrodynamically focus the Chitosan-Dex flow using a sheath flow of ATP-containing water. (b) Effect of flow ratio on hydrodynamic diameter (based on DLS results) and zeta potential of the synthesized NPs. The theoretical mixing time is also shown (filled square

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in blue). (c-e) TEM images of Chitosan-based NPs synthesized at flow ratios of 0.03 (c) and 0.1 (d) in comparison with bulk synthesized NPs (e); (scale bar 100 nm). (f)

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Calculated NP molecular weight (MNP) and local polymer concentration (CNP) inside the Chitosan-based NPs as a function of particle size. (g) Aggregation number (Nagg) of the

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of size. The lines are only guide.

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Chitosan chains in the corresponding NPs and concentration of NPs (NNP) as a function

Figure 2. The effect of NP size and concentration without Dex on MSC viability (based

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on MTT assay).

Figure 3. (a) Effect of NP size, produced using microfluidics platform or by bulk mixing,

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on the process of cellular uptake. The amount of membrane bounded NPs on MSCs and intracellular and total cell uptake of NPs by MSCs are displayed based on CLSM. (b) NP uptake by MSCs based on FACS.

Figure 4. (a) Cumulative in vitro release of Dex from Chitosan/ATP nanoparticles at 37°C and pH of 7.4 (Mean± SD, n>3 independent experiments.) Data are fitted with Eq. 5 (R2>0.98). (b) Calculated diffusion coefficients at 37 °C of Dex within nanoparticles, and relevant time for releasing 50% of loaded drug from the microfluidics-formed and bulk-synthesized nanoparticles.

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ACCEPTED MANUSCRIPT Figure 5. (a) The final structure of Dex drug molecules encapsulated by Chitosan (CS) nanoparticles at equilibrium. CS and Dex atoms are shown in line and stick styles, respectively (carbon: green, oxygen: red, nitrogen: blue, and hydrogen: white); (b)

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Electrostatic, van der Waals and the total interaction energy between CS chains and Dex

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molecules within the last 10 ns of MD simulations. (c) Radial distribution functions of CS amine nitrogen (NH2-CS), hydroxyl oxygen in CS backbone (OH1-CS), and hydroxyl

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oxygen in the hydroxymethyl side chain of CS (OH2-CS) with respect to the double bonded and hydroxyl oxygen in Dex structure; (d) Time evolution of the average number

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of hydrogen bonds between the CS chains and Dex molecules; (e) The average number of hydrogen bonds of -NH2, -OH1, and -OH2 hydroxyl groups in CS encapsulating Dex

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molecules.

Figure 6. Effect of the microfluidics and bulk Dex-loaded Chitosan nanoparticles (NPs)

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on ALP, cbfal, COL1A1 and GAPDH gene expression in mesenchymal stem cells (MSCs). (a) Representative image of RT-PCR. MSCs were cultured in the absence and

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presence of NPs. Representative images of relative mRNA levels of ALP (b), cbfal (c) and COL1A1 (d) in cells were compared and normalized with mRNA of GAPDH in the same

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sample using RT-PCR. MSCs treated with regular complete medium were used as the control sample; MSCs treated with medium containing Dex were used as well; NP-1:

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MSCs treated with medium containing Dex-loaded μF-58 nm NPs; NP-2: MSCs treated with medium containing Dex-loaded μF-74 nm NPs; NP-3: MSCs treated with medium containing Dex-loaded μF-93 nm NPs and NP-4: MSCs treated with medium containing Dex-loaded μF-125 nm NPs; NP-bulk: MSCs treated with medium containing Dexloaded bulk NPs. Values indicate mean ± SD of at least four independent experiments, each performed in triplicate.

Figure 7. ALP activity by protein content (a) without and (b) with Dex. (c) Osteocalcin per DNA content of MSCs after culturing in culture media containing microfluidicsassisted prepared NPs of different sizes for periods of 2, 7 and 14 days. (d) Calcium deposition on extracellular matrix. The cells were transfected with Dex containing CS 27

ACCEPTED MANUSCRIPT microfluidics NPs and cultured for 21 days. Results are expressed as average ± SD, n>4. ##

p < 0.01;

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p < 0.001 between control+Dex and control (without Dex) samples. *p