Flexible Generation of Gradient Electrospinning Nanofibers Using a

May 18, 2012 - The nanofiber surface modified with physical or chemical gradients is very useful in a wide range of areas including tissue engineering...
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Flexible Generation of Gradient Electrospinning Nanofibers Using a Microfluidic Assisted Approach Xu Zhang,†,‡ Xinghua Gao,†,‡ Lei Jiang,† and Jianhua Qin*,† †

Dalian Institute of Chemical Physics, Chinese Academy of Sciences, 457, Zhongshan Road, Dalian, 116023, China Graduate School of the Chinese Academy of Sciences, Beijing, China

‡

ABSTRACT: The nanofiber surface modified with physical or chemical gradients is very useful in a wide range of areas including tissue engineering, regenerative medicine, drug screening, and biomaterial chemistry. In this work, we presented a novel and straightforward microfluidic assisted approach to produce electrospinning nanofibers containing gradients in different compositions, nanoparticles and biomolecule concentrations. The series of gradient nanofibers were mainly produced by using a two inlet microfluidic device in combination with an electrospinning nozzle on a 3-D controllable platform, which exhibited different functions and properties. The controlled nanofibers with incorporated biomolecule gradient were used for guiding the spatial differentiation in mesenchymal stem cells (MSCs). This established approach is very simple, and flexible to operate, which might find enormous potential for biology and tissue engineering applications.

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INTRODUCTION Electrospinning is a simple and versatile technique to generate nanoscale fibers, which has attracted much interest in a broad range of polymer systems.1 Due to the characteristics of porosity, large specific surface area, wide selection of material, and bioarchitecture suitability, electrospinning nanofibers have been utilized as biocompatible scaffolds or drug/DNA carrier for use in the fields of tissue engineering, regenerative medicine, and drug delivery. Presently, several works have explored functionalization of the electrospun nanofibers by modifying their surfaces, thus enhancing the performance for extended applications. For example, the topology of electrospinning nanofibers can be modified to super porous or grid-shaped to maintain the functions of neurocytes and cardiac cells.1−3 Nanofibers can also be coated with proteins in order to form a special matrix for cell culture,4,5 and modified with drugs/DNA to guide the cell growth, apoptosis, and differentiation.6−11 Furthermore, it can be functionalized with special architecture, drugs, or nanoparticles to repair or rebuild tissues, such as bone, vein, and skin.1,12−18 However, common electrospinning methods can only generate nanofibers with the same concentration, which is not sufficient for mimicking the complex environment in vivo that consists of different kinds of matrix, physical, and © 2012 American Chemical Society

chemical signal gradients. Materials bearing favorable functionalities (hydrophilic, wettability, biomolecule modified, and so forth) are generally needed for interfacial tissue engineering. Recently, some researchers reported a simple method to produce electrospinning nanofibers with surface-incorporated gradients using a controlled wet filling method.19,20 Using this method, one side of the nanofibers was immersed in a solution containing functional molecules; then, the gradient could be generated along the axis of the fiber surface under capillary force. However, this method still has some shortcomings which might potentially limit its practical use. The wetting process might change the inherent properties of the nanofibers; thus, it is not suitable for producing the gradients containing nanoparticles and components with enhanced functionality. Particularly, this method simply coated the functional molecules on the surfaces of the fibers rather than inserting into the internal fibers, which might sacrifice the advantages of the gradient nanofibers and influence the release efficiency of the functional fibers. Received: February 26, 2012 Revised: April 30, 2012 Published: May 18, 2012 10026

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In order to overcome the current limitations and generate nanofibers with controlled gradients and enhanced functionality, we presented a novel and flexible approach to produce a series of nanofiber gradients via electrospnning mixed solutions by combining a microfluidic device with an electrospining nozzle on a 3-D moving platform. Using this approach, the different proportions of polymer and particle solutions could be mixed and electrospun flexibly, thus producing the spatially controlled gradients with different properties. In particular, the gradient nanofibers could be generated within a controllable spun area and collected continuously, which is not possible by other method. On the basis of this approach, the three types of gradient nanofibers incorporated with different polymer components, particles, and biomolecule concentrations were demonstrated. Furthermore, the fibers modified with biomolecule gradient were presented for biocompatible cell culture and guided differentiations in mesenchymal stem cells (MSCs).

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EXPERIMENTAL SECTION Figure 1. Schematic diagram (A) and photograph (B) of the platform to generate nanofibers with flexible gradient.

Materials. Polydimethylsiloxane (PDMS) (Sylgard 184, Dow Corning, Midland, MI) was purchased from Dow Corning company. SU-8 (3035) photoresist was purchased from MicroChem. Poly(vinylpyrrolidone) (PVP: 360 000 Da) was purchased from MP Biomedical LLC. Poly(lactic-co-glycolic acid) (PLGA: 1.64 dL/g) was purchased from Changchun Sibobiomaterials company. Gelatin (G9382) was purchased from Sigma. Trifluoroethanol was purchased from Aladdin Chemistry Company. IgG-FITC and Rhodamine B were purchased from Zhongshan Goldenbridge Bio. Company. Teflon (PTFE Beads microdispers-200, 80 000 Da, mean particle size: 200− 300 nm, surface area: 10 m2/g) was purchased from Polysciences, Inc. Dexamethasone was purchased from Sigma. AB cement was purchased from ALTECO Company (F-05). Red and blue dyes were color pen inks 20% in ethanol (volume %). Ethanol (analytically pure 99.7%) and deionized water were used for experiments. Solutions Prepared. PVP was dissolved in ethanol with concentrations of 8%, 12%, and 16% (w/v, unit g/mL) by stirring for 10 min. PLGA was dissolved in trifluoroethanol with concentrations of 8%, 12%, 16%, and 20% (w/v, unit g/mL) by stirring for 2−4 h. Gelatin was dissolved in trifluoroethanol with concentrations of 10%, 15%, and 20% (w/v, unit g/mL) by stirring for 6−8 h. Rhodamine B and IgG-FITC were added into 12% PVP/ ethanol solution directly followed by mixing for 5 min. Dexamethasone was added into 12% PLGA/trifluoroethanol solution directly followed by mixing for 5 min. 40% PTFE beads (w/v, unit g/mL) were dispersed into 5% PLGA/trifluoroethanol solution by ultrasonic treatment for 6−8 h, in which the PLGA fragments would be attached onto the PTFE surface to increase their dispersibility. Then, this solution was mixed with 20% PLGA/trifluoroethanol solution to generate 20% PTFE/12%PLGA trifluoroethanol solution. Fabrication of Microfluidic Device. The Y-shaped designed microfluidic device was fabricated using the soft-lithography method, which was used as a mixer for electrospun polymer solutions. As shown in Figure 1, the device consisted of a PDMS block featured with microchannels (500 μm wide and 150 μm deep) and a slide glass, which were irreversibly sealed by oxygen plasma for 1 min. To connect with high voltage for electrospinning, a metal tube (Φex 500 μm, Φin 250 μm) with a thinner nozzle (Φex 225 μm, Φin 100 μm) was inserted into the outlet of the microfluidic device and airproofed by AB cement. Control of Electrospun Area Size. In this work, a 3-D moving platform was utilized to control the electrospining distance between a collector and a nozzle, and this was able to control the size of the electrospun area. The distance between the electrospinning nozzle and the collector was 1.5 cm and the applied voltage was 3 kV ± 0.3 kV. In this way, the diameters of the spinning areas could be decreased to 3− 5 mm. (3-D moving platform was homemade using 3 stepping motors purchased from ZoLix. Company, PSA100−11AS-X, step sizes 0.3125 μm, ranges 100 mm, speeds possible 3.125 mm/s.)

Operation of Gradient Generation. As shown in Figure 1, the microfluidic chip has a Y-shaped channel including an outlet and two inlets. In this experiment, the two inlets were connected with two syringes (sterile for single use, volume 1 mL) that were controlled by two pumps (purchased from LongerPump Company, TJ-1A), respectively. One type of polymer solution containing functional molecules is defined as 100%, and the other polymer solution without functional molecules is defined as 0%. The two polymer solutions were injected from two syringes and completely mixed in a serpentine microchannel, and the different concentrations of function molecules in polymers were obtained by adjusting the volumetric ratio (syringe flow rates) of the two polymer solutions at the inlets. Furthermore, using a 3D moving platform, the nanofibers with different concentrations could be accurately spun within the spinning area to form concentration gradients. The residues of the solvents were removed from the nanofibers by using a vacuum oven for 8−12 h. All the proportions of syringe flow rates and functional molecule concentrations used in this work are described in Tables 1−4.

Table 1. Proportions of Polymers to Generate Nanofibers with Different Componentsa solutions flow rate

16% gelatin (w/v)

12% PLGA (w/v)

G/P proportion (v/v)

1 2 3 4 5

80 μL/h 60 μL/h 40 μL/h 20 μL/h 0 μL/h

0 μL/h 20 μL/h 40 μL/h 60 μL/h 80 μL/h

4:0 3:1 2:2 1:3 0:4

a

w/v means mass/volume, unit g/mL, and v/v means volume/volume.

Table 2. Proportions of Teflon Nanoparticle (PTFE) Concentrations in PLGA Nanofibers (a) solutions

a

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flow rate

12% PLGA (w/v) + 20% PTFE (w/v)

12% PLGA (w/v)

PTFE/mixed solution (w/v)

1 2 3 4

0 μL/h 20 μL/h 40 μL/h 80 μL/h

80 μL/h 60 μL/h 40 μL/h 0 μL/h

0% 5% 10% 20%

w/v means mass/volume, unit g/mL. dx.doi.org/10.1021/la300821r | Langmuir 2012, 28, 10026−10032

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solution at a certain flow rate. The time required to finish the replacement process was calculated as delay time. In order to measure the delay time, three kinds of polymer solutions with different concentrations/viscosities (PVP, 8%, 12%, 16%; PLGA, 8%, 12%, 16%; Gelatin, 10%, 15%, 20%) were injected into the channel at varied flow rates (40 μL/min, 60 μL/min, 80 μL/min, 100 μL/min) (Figure 2B). Cell Culture. The MSCs were derived from mouse bone marrow, which was supplied by Department of Cell Biology, Institute of Basic Medical Sciences, China, as a gift.21 The cells were cultured in α-MEM containing 15% fetal bovine modified serum (FBS; HyClone, USA), 100 μg/mL streptomycin, and 100 U/mL penicillin, at 37 °C in a wetting atmosphere containing 95% air and 5% CO2, and the cells was passaged every 3 days. For differentiation experiments on the nanofibers gradient, the nanofiber scaffold was first modified with dexamethasone gradient (10−3 mol/L, 6 × 10−4 mol/L, 4 × 10−4 mol/ L, 2 × 10−4 mol/L). Then, a PDMS block with 12 (3 × 4) small wells (Φ 2.5 mm) was covered on the nanofibers. The seeded MSCs in the wells were cultured for 14 days. The medium was α-MEM containing 15% fetal bovine serum and supplemented with 5 × 10−5 mol/L ascorbic acid and 10−2 mol/L β-glycerol phosphate (Sigma-Aldrich), which was added every 3 days.

Table 3. Proportions of Rhodamine B and Human IgG-FITC in PVP Nanofibersa A solutions flow rate 1 2 3 4 5

12% PVP (w/v) + 5 × 10−10 g/L RhB (defined as 100%) 16 32 48 64 80

12% PVP (w/v) (defined as 0%)

RhB relative concentration

64 μL/h 48 μL/h 32 μL/h 16 μL/h 0 μL/h

20% 40% 60% 80% 100%

μL/h μL/h μL/h μL/h μL/h B solutions

flow rate

12% PVP (w/v) + 1/30 IgG (v/v) (defined as 100%)

12% PVP (w/v) (defined as 0%)

IgG relative concentration

1 2 3 4 5

0 μL/h 20 μL/h 40 μL/h 60 μL/h 80 μL/h

80 μL/h 60 μL/h 40 μL/h 20 μL/h 0 μL/h

0% 25% 50% 75% 100%

a

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w/v means mass/volume, unit g/mL, and v/v means volume/volume.

Table 4. Proportions of Dexamethasone in PLGA Nanofibersa solutions flow rate 1 2 3 4 a

12% PLGA (w/v) + 10−4 mol/ LDex (defined as 100%) 16 32 48 80

12% PLGA (w/v) (defined as 0%)

Dex relative concentration

64 μL/h 48 μL/h 32 μL/h 0 μL/h

20% 40% 60% 100%

μL/h μL/h μL/h μL/h

RESULTS AND DISCUSSIONS

Optimization of the Parameters to Generate Gradient Electrospun Nanofibers. In this work, the gradient nanofibers were generated by combining a microfluidic device with an electrospinning nozzle on a 3-D moving platform. In order to produce the gradient fibers accurately, it is necessary to optimize several critical parameters involved in the experiment, including the uniformity and the delay time of electrospun polymer solutions in the mixing process, electrospinning distance, voltage, flow rate, and electrospun area. Prior to electrospinning process, the two types of polymer solutions were initially mixed by a two-inlet microfluidic device driven by pumps and then electrospun from a nozzle. Under various flow rates, the proportions of the two solutions could be perfectly controlled by computer-programmed pumps, thus leading to the production of nanofibers with different components. Specially, by using a 3-D moving platform on the bottom, any position within the spun area could be located and the gradient of nanofibers on the collector could be controlled flexibly.

w/v means mass/volume, unit g/mL.

The delay time of the device is an important parameter to influence the accuracy of gradient generation. Here, the delay time indicates the interval between the starting point of solutions mixing in the microchannel to the end point of mixing solution reaching the electrospinning nozzle. This delay time was measured by a solution replacement method, as shown in Figure 2A. At first, the channel was filled with a red solution, followed by the replacement of a blue

Figure 2. Condition optimization of the microfluidic device. (A) Schematic diagram of the solution replacement method to test the delay time. (B) Delay times of three polymers plotted versus flow rates. 10028

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Figure 3. Generating nanofibers with gelatin and PLGA compound gradient. (A) Schematic diagram of the mixing process in a microfluidic chip. (B) SEM images of nanofibers produced by different proportion of gelatin and PLGA (4:0, 3:1, 2:2, 1:3, 0:4) and the proportion of carbon and oxygen in these fibers.

the normal one by increasing the electrospinning distance, voltage, and flow rate. Nanofiber Gradients Incorporated with Different Polymer Components. On the basis of the optimized conditions, this method can be further used to generate nanofibers with physical and chemical gradients. At first, gradient nanofibers with two types of component were produced, in which the component proportions were varied. In this way, the properties of nanofibers could be changed according to the properties of the electrospun polymers. Initially, PLGA and gelatin were chosen to generate compound gradient nanofibers, because they could be dissolved in the same solvent as trifluoroethanol. As shown in Table 1, 16% gelatin and 12% PLGA were dissolved into trifluoroethanol and proportionally mixed in the microchannel (4:0, 3:1, 2:2, 1:3, 0:4 v/v), followed by electrospinning. The SEM results showed that high proportions of gelatin (4:0, 3:1 v/v) may produce electrospinning fibers with large diameters (more than 1 μm), and the space between the fibers was connected by thin membranes. In contrast, the diameters of the fibers with high proportions of PLGA were smaller (500−600 nm), and no polymer membrane was observed between fiber intervals (in Figure 3). These different results were mainly dependent on the different molecule weight and space structure of the polymers. Through X-ray scanning, the major elements including carbon and oxygen were seen in the fibers, in which the concentrations were varied with the different proportions of polymers. This means that the ingredients of fibers were changed and they could be modified inside, not limited to the change on the fiber surface. The component-gradient nanofiber scaffold has the potential to satisfy the requirement of tissue healing, involving different kinds of matrix, such as bone, cartilage, connective tissue, and muscle. Furthermore, the surface of the ingredient-gradient nanofibers can be used in material science and surface chemistry applications. Nanofiber Gradient Incorported with Nanoparticles. In addition to generated nanofibers with different component gradients, this method can be used to generate nanofibers with different nanoparticle gradients as well. This gradient can offer nanofibers with different physical properties including roughness, hardness, and hydrophilic−hydrophobic gradients.

During the mixing and electrospinning process, the delay time exists, which indicates the time between the starting point of solutions mixing in microchannel and the end point of mixed solution arriving in the nozzle. This value was very important for maintaining the stability of this system and controlling the fiber gradient accurately. The delay time was calculated by the solution replacement method as depicted in Figure 2A. As shown in Figure 2B, the polymer solutions had a similar changing trend in their delay time. Both the pump flow rate and the polymer concentration (solution viscosity) influence the delay time. In solutions with a low polymer concentration, such as 8% and 12%, the delay times declined obviously with the increasing of flow rate (40 μL/h to 100 μL/h). However, in high-concentration solutions (16% and 20%), the delay time was not changed with the flow rate, especially at high flow rate (greater than 80 μL/h). The phenomenon was mainly due to the back pressure produced in the channel. Thus, in this work, the optimum flow rate (80 μL/h) and polymer concentration (12% PLGA, 12% PVP, 16% Gelatin) were selected in the following work. The electrospinning distance is another major factor to be considered, because it is positively related to the size of the electrospun area. In this work, a 3-D moving platform was utilized to control the electrospining distance between a collector and the nozzle. It was observed that, when the electrospinning distance was reduced from 10 cm to 1 cm, the diameter of electrospun area could be decreased from 30−50 mm to 2−3 mm. Thus, the applied voltage and pump flow rate should be reduced from 10 to 2 kV and from 500 to 60 μL/h, accordingly. It is observed that this approach is quite convenient for producing electrospinning nanofibers on a very small coverslip (2.3 cm × 2.3 cm), which is not possible using the common method. The common electrospinning operation usually generates a relatively large electrospun area, thus limiting its ability to generate gradient fibers. Also, the large area scale of fibers is not compatible with the small size of plate wells or culture dishes in microscale cellular assays. Here, under the optimized electrospinning distance (1.5 cm), pump flow rate (80 μL/h), and electrospinning voltage (3 ± 0.3 kV), the electrospun area with a diameter of 3−5 mm can be produced. Certainly, by using the 3-D moving platform, the size of the electrospinning area formed could easily be as large as 10029

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Figure 4. Generated nanofibers with nanoparticle gradient to change surface properties. (A) SEM images of nanofibers with different nanoparticle concentrations and the photos of droplets on them. (B) Top view and side view of droplets on nanofiber surface with continuous nanoparticle gradient. (C) Droplet contact angles on different nanofiber surfaces with different nanoparticle concentrations.

Here, Teflon nanoparticles (PTFE) were used as a hydrophobic material to generate a hydrophilic−hydrophobic gradient in nanofibers. PLGA nanofibers with four Teflon nanoparticle concentrations were produced, including 0%, 5%, 10%, and 20% (w/v) (Table 2). The nanoparticles with different concentrations could be mixed evenly into the fibers, but the conglomerate size of nanoparticles across the fiber surfaces increased with increasing particle concentrations, indicating that the dispersibility of nanoparticles is influenced by particle concentrations. Also, the same-sized droplets were dipped on the nanofiber surfaces to reflect the wetting property of the surface (Figure 4A). The series of droplets with different contact angles represented the formation of a hydrophilic− hydrophobic gradient, as shown in Figure 4B. The contact angles of droplets increased with the increase of nanoparticle concentrations within nanofibers. The range of contact angles could span from 104° to 142° between minimum and maximum concentrations, as shown in Figure 4C. The results demonstrated that the gradient wettability property of the nanofibers could be demonstrated in a controllable manner. Actually, this physical property of surfaces can guide cell behavior, such as cell migration and differentiation.22−24 In particular, the hydrophobic gradient surface also has the potential to be used to drive the microflow relying on different surface tensions.25 Nanofiber Incorporated Biomolecule Gradient for Guided Spatial Differentiation in MSCs. The established approach can further be used to generate nanofibers with biochemical molecule gradients, such as proteins and drugs. To test the feasibility of this method, two fluorescence indicators with different molecule weights were used to characterize the functional gradients in nanofibers, including Rhodamine B (479 Da) and human IgG-FITC (160 000 Da). The concentrations of these two indicators were listed in Table 3 and the fluorescence intensity within fibers would reflect their actual concentration. In Figure 5, the fluorescence intensity in nanofibers was plotted versus relative concentration, and both working curves of the Rhodamine B and IgG-FITC have good linear relationships (R = 0.9996 and 0.9951), which indicated the molecules with different molecule weights could be mixed into the fibers proportionally.

Figure 5. Characterization of biochemical gradient in nanofibers by using Human IgG-FITC and Rhodamine B. The work curves were fluorescence intensity of nanofibers plotted versus relative concentration of Human IgG and Rhodamine B in the fibers.

The unique property of this method is that it can generate biomolecule concentration gradients without the requirement of wetting process. For the protein gradient, if the proteins may be denatured in organic solution, the problem can be solved by electrospinning emulsified proteins.26 We believe that this scaffold with biochemical molecule gradients may be extended to use in 3-D cell culture with different concentrations of stimulus or localized drug release. MSCs is a type of multipotent cell, which can be proliferated in vitro, and has the potential to differentiate into lines of mesenchymal tissues, such as bone, cartilage, fat, and so on. In the following work, a nanofiber scaffold with a dexamethasone concentration gradient was produced to spatially induce differentiation in MSCs. In traditional experiments, MCSs could be induced into adipocytes by using induction medium with 10−6 mol/L dexamethasone or induced into osteocytes by using induction medium with 10−7 mol/L dexamethasone.21,27 In this work, four dexamethasone concentrations (10−4 mol/L, 6 × 10−5 mol/L, 4 × 10−5 mol/L, 2 × 10−5 mol/L) formed PLGA nanofibers that were electrospun on a cover slide continuously, and the proportions were defined as 100%, 60%, 10030

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Figure 6. Nanofibers with dexamethasone concentration gradient induce MSC differentiation. (A) Fluorescent images of MSC attachment and proliferation grown on dexamethasone gradient nanofibers. (B) Schematic diagram of MSC specific differentiation induced by substrate (i). ALP (osteocyte) and red-oil (adipocyte) staining images of MSCs growth on nanofibers with dexamethasone concentration gradient (ii). Differentiation proportion of MSCs induced by substrate with different dexamethasone concentrations (iii).

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CONCLUSIONS This work demonstrated a novel and flexible approach to generate gradient nanofibers by using a microfluidic device in combination with an electrospinning nozzle. This approach enables the generation of spatially controlled fibers incorporated with different polymer compositions, nanoparticles, and biomolecule concentrations, which exhibit enhanced functionalities. The fibers modified with biomolecule gradient demonstrated a strong ability to spatially direct the differentiation in MSCs. This method can generate gradient nanofibers with different incorporating fashions, which will supply complicated nanofibers with varied chemical and physical properties. We believe this approach is an advance over previous methods due to its versatility and controllability, indicating the enormous potential in tissue engineering, regenerative medicine, and material science. Particularly, it can be extended to produce more complex gradient nanofibers with improved functionality by introducing more microfluidic elements to the device.

40%, and 20% (dexamethasone relative concentration) in Table 4. A PDMS block was covered on the nanofiber surface with small wells (Φ 2.5 mm) to restrict the area for cell growth. The MSCs were cultured on the surface restricted by the wells. As shown in Figure 6A, the adhesion and proliferation of MSCs grown on nanofibers with dexamethasone gradients were as good as that of cells grown on a glass slide. However, the cells on fibers exhibited a typical 3-D morphology which is close to mimicking the in vivo physiologically relevant microenvironment. After 14 days culture, MSCs could be differentiated into specific kinds of cells at localized areas on the same surface (Figure 6B). The results showed that MSCs were induced into adipocytes in the area of fibers with higher drug concentration and induced into osteocytes in the area of fibers with lower drug concentration (Figure 6B ii). Furthermore, the proportion of dexamethasone was positively related to the adipogenic differentiation (R = 0.9953) but negatively related to the osteogenic differentiation (R = 0.9995). These results demonstrated that MSC differentiation could be spatially controlled using this scaffold, which supplied a potential tool for interfacial tissue engineering. For example, when rebuilding a tissue in vitro, it is very difficult to distribute different kinds of cells into specific space according to the tissue structure. This nanofiber scaffold with biochemical molecule gradients can guide stem cell differentiation with spatial control properties, thus providing a unique tool to overcome the difficulty. Of course, this method could be further improved by mixing more kinds of biochemical factors into nanofibers and applying more complex stimulus in tissue engineering applications with enhanced functionality.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by Joint Research Fund of NSFCRGC (11161160522), Knowledge Innovation Program of the Chinese Academy of Sciences (KJCX2-YW-H18), and Instru10031

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Scaffolds for Tissue Engineering. ACS Appl. Mater. Interfaces 2010, 2 (4), 1025−1030. (20) Li, X. R.; Xie, J. W.; Lipner, J.; Yuan, X. Y.; Thomopoulos, S.; Xia, Y. N. Nanofiber scaffolds with gradations in mineral content for mimicking the tendon-to-bone insertion site. Nano Lett. 2009, 9 (7), 2763−2768. (21) Gao, J.; Yan, X.-L.; Li, R.; Liu, Y.; He, W.; Sun, S.; Zhang, Y.; Liu, B.; Xiong, J.; Mao, N. Characterization of OP9 as authentic mesenchymal stem cell line. Journal of Genetics and Genomics 2010, 37 (7), 475−482. (22) Engler, A. J.; Sen, S.; Sweeney, H. L.; Discher, D. E. Matrix Elasticity Directs Stem Cell Lineage Specification. Cell 2006, 126 (4), 677−689. (23) Tse, J. R. ; Engler, A. J. Stiffness Gradients Mimicking In Vivo Tissue Variation Regulate Mesenchymal Stem Cell Fate. PLOS One 2011, 6 (1), e15978. (24) Lo, C.-M.; W., H.-B.; Dembo, M.; Wang, Y.-L. Cell Movement Is Guided by the Rigidity of the Substrate. Biophys. J. 2000, 79 (1), 144−152. (25) Xing, S.; Harake, R. S.; Pan, T. Droplet-driven transports on superhydrophobic-patterned surface microfluidics. Lab Chip 2011, 11 (21), 3642−3648. (26) Yang, Y.; Xia, T.; Zhi, W.; Wei, L.; Weng, J.; Zhang, C.; Li, X. Promotion of skin regeneration in diabetic rats by electrospun coresheath fibers loaded with basic fibroblast growth factor. Biomaterials 2011, 32 (18), 4243−4254. (27) Gao, X. H.; Zhang, X.; Tong, H. Y.; Lin, B. C.; Qin, J. H. A simple elastic membrane-based microfluidic chip for the proliferation and differentiation of mesenchymal stem cells (MSCs) under tensile stress. Electrophoresis 2011, 32 (23), 3431−3436.

ment Research and Development Program of the Chinese Academy of Sciences (YZ200908).

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REFERENCES

(1) Xie, J.; MacEwan, M. R.; Schwartz, A. G.; Xia, Y. Electrospun nanofibers for neural tissue engineering. Nanoscale 2010, 2 (1), 35. (2) Gulfam, M.; Lee, J. M.; Kim, J.-e.; Lim, D. W.; Lee, E. K.; Chung, B. G. Highly Porous Core−Shell Polymeric Fiber Network. Langmuir 2011, 27 (17), 10993−10999. (3) Orlova, Y.; Magome, N.; Liu, L.; Chen, Y.; Agladze, K. Electrospun nanofibers as a tool for architecture control in engineered cardiac tissue. Biomaterials 2011, 32 (24), 5615−5624. (4) Ulrich, T. A.; Jain, A.; Tanner, K.; MacKay, J. L.; Kumar, S. Probing cellular mechanobiology in three-dimensional culture with collagen−agarose matrices. Biomaterials 2010, 31 (7), 1875−1884. (5) Hashemi, S. M.; Soudi, S.; Shabani, I.; Naderi, M.; Soleimani, M. The promotion of stemness and pluripotency following feeder-free culture of embryonic stem cells on collagen-grafted 3-dimensional nanofibrous scaffold. Biomaterials 2011, 32 (30), 7363−7374. (6) Nie, H.; Ho, M.-L.; Wang, C.-K.; Wang, C.-H.; Fu, Y.-C. BMP-2 plasmid loaded PLGA/HAp composite scaffolds for treatment of bone defects in nude mice. Biomaterials 2009, 30 (5), 892−901. (7) Nie, H.; Wang, C.-H. Fabrication and characterization of PLGA/ HAp composite scaffolds for delivery of BMP-2 plasmid DNA. J. Controlled Release 2007, 120 (1−2), 111−121. (8) Fu, S. Z.; Wang, X. H.; Guo, G.; Shi, S.; Liang, H.; Luo, F.; Wei, Y. Q.; Qian, Z. Y * Preparation and Characterization of Nanohydroxyapatite/poly(ε-caprolactone)- poly(ethyleneglycol)-poly(εcaprolactone) composite fibers for tissue engineering. J. Chem. Phys. 2010, 114 (43), 18372−18378. (9) Gupta, D.; Venugopal, J.; Mitra, S.; Giri Dev, V. R.; Ramakrishna, S. Nanostructured biocomposite substrates by electrospinning and electrospraying for the mineralization of osteoblasts. Biomaterials 2009, 30 (11), 2085−2094. (10) Ignatious, F.; Sun, L.; Lee, C.-P.; Baldoni, J. Electrospun nanofibers in oral drug delivery. Pharm. Res. 2010, 27 (4), 576−588. (11) Tiwari, S. K.; Tzezana, R.; Zussman, E.; Venkatraman, S. S. Optimizing partition-controlled drug release from electrospun core− shell fibers. Int. J. Pharm. 2010, 392 (1−2), 209−217. (12) Pişkin, E.; Işoğlu, I. A.; Bölgen, N.; Vargel, I. b.; Griffiths, S.; Ç avuşoğlu, T.; Korkusuz, P.; Güzel, E.; Cartmell, S. In vivo performance of simvastatin-loaded electrospun spiral-wound polycaprolactone scaffolds in reconstruction of cranial bone defects in the rat model. J. Biomed. Mater. Res., Part A 2009, 90A (4), 1137−1151. (13) Hashi, C. K.; Zhu, Y.; Yang, G. Y.; Young, W. L.; Hsiao, B. S.; Wang, K.; Chu, B.; Li, S. Antithrombogenic property of bone marrow mesenchymal stem cells in nanofibrous vascular grafts. Proc. Natl. Acad. Sci. U.S.A. 2007, 104 (29), 11915−11920. (14) Yang, X. C.; Sc, M.; Shah, J. D.; M., E.; Wang, H. J. Nanofiber enabled layer-by-layer approach toward three-dimensional tissue formation. Tiss. Eng. 2009, 15 (4), 1945−1956. (15) K-hasuwan, P.-r.; Pavasant, P.; Supaphol, P. Effect of the Surface Topography of Electrospun Poly(ε-caprolactone)/Poly(3-hydroxybuterate-co- 3-hydroxyvalerate) Fibrous Substrates on Cultured Bone Cell Behavior. Langmuir 2011, 27 (17), 10938−10946. (16) Nandakumar, A.; Yang, L.; Habibovic, P.; van Blitterswijk, C. Calcium Phosphate Coated Electrospun Fiber Matrices as Scaffolds for Bone Tissue Engineering. Langmuir 2010, 26 (10), 7380−7387. (17) Li, X. R.; Xie, J. W.; Yuan, X. Y.; Xia, Y. N. Coating Electrospun Poly(ε-caprolactone) Fibers with Gelatin and Calcium Phosphate and Their Use as Biomimetic Scaffolds for Bone Tissue Engineering. Langmuir 2008, 24 (24), 14145−14150. (18) Lewkowitz-Shpuntoff, H. M.; Wen, M. C.; Singh, A.; Brenner, N.; Gambino, R.; Pernodet, N.; Isseroff, R.; Rafailovich, M.; Sokolov, J. The effect of organo clay and adsorbed FeO3 nanoparticles on cells cultured on Ethylene Vinyl Acetate substrates and fibers. Biomaterials 2009, 30 (1), 8−18. (19) Shi, J.; Wang, L.; Zhang, F.; Li, H.; Lei, L.; Liu, L.; Chen, Y. Incorporating Protein Gradient into Electrospun Nanofibers As 10032

dx.doi.org/10.1021/la300821r | Langmuir 2012, 28, 10026−10032