cold atmospheric plasma (cap)-modified and bioactive...
TRANSCRIPT
COLD ATMOSPHERIC PLASMA (CAP)-MODIFIED AND BIOACTIVE PROTEIN-LOADED CORE-SHELL NANOFIBERS FOR BONE TISSUE ENGINEERING
APPLICATIONS
A Thesis Presented
By
Yangfang Zhou
to
The Department of Chemical Engineering
in partial fulfillment of the requirements for the degree of
Master of Science
in the field of
Chemical Engineering
Northeastern University Boston, Massachusetts
May, 2018
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ABSTRACT Coaxial electrospinning is a novel technique for producing core-shell nanofibers that
combine the advantages of having a robust structure and the ability to deliver hydrophilic
bioactive agents. The objective of the study was to fabricate core-shell nanofibers
comprised scaffolds that can not only simulate physiological extracellular matrix (ECM),
but can also deliver bioactive proteins in a controlled manner. It was hypothesized that the
size of scaffold surface pores could be increased by cold atmospheric plasma (CAP)
modification. By modifying the surface pore size, release profiles of loaded agents could
subsequently be controlled. To evaluate this hypothesis, bovine serum albumin (BSA) was
used as a model protein to be loaded into the core polymer. After optimization of coaxial
electrospinning parameters, uniform and bead-free polyvinyl alcohol (PVA)/poly (L-lactic
acid) (PLLA) core-shell nanofibers were fabricated. Additionally, CAP was used to create
nano-structured surfaces, increase surface pore size, and change surface hydrophilicity of
electrospun scaffolds. After CAP treatment, water contact angles were reduced from 110°
to 50° and protein and water adsorption was significantly elevated. The changes in
hydrophilicity and improved scaffold surface area contributed to a dramatic promotion of
fibroblasts and osteoblasts attachment and proliferation. Also, alkaline phosphatase (ALP)
activity, total protein content and calcium deposition from mesenchymal stem cells (MSCs)
on the scaffolds showed that the CAP-modified nanofibrous scaffolds were more
osteoinductive. Most importantly, CAP treatment resulted in an increase in nanofiber
surface pore size, which can further contribute to significant changes in drug release
profiles. FITC-dextran was used as a fluorescent probe in drug release study. After 120
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hours, the CAP-treated samples released almost 100% of the loaded FITC-dextran, while
the untreated samples only released 40% of the loaded FITC-dextran. Collectively, these
results indicate that CAP-treated and bioactive protein-loaded core-shell nanofibers can be
valuable in regenerative medicine next generation as a drug delivery system in bone tissue
engineering.
Keywords: Coaxial electrospinning, Cold atmospheric plasma (CAP), Surface
modification, Drug delivery, Bone tissue engineering.
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ACKNOWLEDGMENTS
I would like to thank Northeastern University for supporting my research and every
faculty member and staff in the Department of Chemical Engineering.
Firstly, I want to thank Professor Thomas J. Webster who gave me the opportunity to
conduct my research in his amazing Nanomedicine Lab. He is always willing to provide
innovative and helpful advice on my research. Additionally, he always supports and
encourages us to attend different conferences, which are excellent opportunities for us to
enhance professional skills. Moreover, I would like to thank Dr. Mian Wang who brought
me into the research of electrospun nanofibers. He was my coworker as well as my mentor
who trained me various experimental and writing skills. Also, I would like to thank other
lab members who are so friendly and helpful. Furthermore, I would like to thank William
Fowle who trained me to use the SEM and Robert R. Eagan who helped me print posters
and repaired the electrospinner in our lab. I also would like to thank Ian Harding who
helped me revise the manuscript and gave me numerous helpful suggestions. I appreciate
the help and advice from my thesis committee members: Professor Ambika Bajpayee and
Professor Arthur Coury. Last, I would like to thank my parents who raised me up in such
a happy family and supported me to study in the United States.
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TABLE OF CONTENTS
1. Introduction ................................................................................................................. 1
1.1. Motivation and Background ............................................................................... 1
1.2. Objectives and Scope .......................................................................................... 2
2. Critical Literature Review ........................................................................................... 4
2.1. Bone Grafts and Bone Tissue Engineering (BTE) .............................................. 4
2.2. The Extracellular Matrix (ECM) of Bone ........................................................... 6
2.3. Nano-fibrous Scaffolds in Bone Tissue Engineering .......................................... 8
2.4. Electrospinning ................................................................................................... 9
2.5. Poly (L-lactic acid) (PLLA) .............................................................................. 11
2.6. Polyvinyl Alcohol (PVA) ................................................................................. 12
2.7. Cold Atmospheric Plasma (CAP) ..................................................................... 12
2.8. Summary ........................................................................................................... 14
3. Materials and Methods .............................................................................................. 14
3.1. Optimization of Coaxial Electrospinning Parameters ....................................... 14
3.2. CAP Treatment ................................................................................................. 16
3.3. Evaluation of Surface Morphology ................................................................... 17
3.4. Measurements of Surface Wettability ............................................................... 17
3.5. In vitro Drug Release ........................................................................................ 17
3.6. Osteoblast and Fibroblast Response on Nanofibrous Scaffolds ....................... 18
3.6.1. Cell Culture ............................................................................................... 18
3.6.2. Cell Proliferation Study ............................................................................ 18
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3.6.3. Osteoblast Vinculin Staining .................................................................... 19
3.7. Mesenchymal Stem Cell (MSC) Response on Nanofibrous Scaffolds .............. 20
3.7.1. MSC Culture ............................................................................................. 20
3.7.2. MSC Differentiation ................................................................................. 20
3.7.3. ALP Activity Measurement ...................................................................... 20
3.7.4. Total Protein Content Measurement ......................................................... 21
3.7.5. Calcium Content Measurement ................................................................. 21
3.8. Protein Adsorption ............................................................................................ 21
3.9. Analysis of Scaffold Swelling and Water Content ........................................... 22
3.10. Statistical Analysis ............................................................................................ 23
4. Results ....................................................................................................................... 23
4.1. Surface Morphology of Nanofibers .................................................................. 23
4.2. In vitro Drug Release ........................................................................................ 27
4.3. Water Contact Angle ......................................................................................... 28
4.4. Osteoblast and Fibroblast Response on Nanofibrous Scaffolds ....................... 29
4.5. MSC Response on Nanofibrous Scaffolds ........................................................ 32
4.6. Protein Adsorption and Water Swelling ........................................................... 35
5. Discussion ................................................................................................................. 37
5.1. Optimization of Coaxial Electrospinning Parameters ....................................... 37
5.2. Cell Behavior .................................................................................................... 38
6. Conclusions ............................................................................................................... 40
7. Future Directions ...................................................................................................... 41
7.1. Bone Tissue Engineering .................................................................................. 41
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7.2. Wound Healing ................................................................................................. 41
REFERENCES ................................................................................................................. 43
APPENDIX ....................................................................................................................... 53
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LIST OF TABLES Table 1. Advantages and disadvantages of three types of bone grafts (modified from7). .. 4
Table 2. Optimization of coaxial electrospinning parameters. ......................................... 16
Table 3. Morphology of nanofibers fabricated by coaxial electrospinning with different
parameters. ................................................................................................................ 25
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LIST OF FIGURES Figure 1. Paradigm of bone tissue engineering (adopted from 8). ...................................... 6
Figure 2. Schematic of the extracellular matrix (adopted from 12). .................................... 7
Figure 3. Schematic of current fabrication techniques of nano-fibrous scaffolds. Scale bar:
10 µm (a, c), 500 nm (b) (adopted from 12). ............................................................... 9
Figure 4. Schematic diagram of the coaxial electrospinning (modified from 16). ............ 10
Figure 5. Schematic of the cold atmospheric plasma setup (adopted form 47, 48). ............ 13
Figure 6. Possible mechanisms of chemical and roughness changes caused by CAP on PLA
scaffolds (adopted form 33). ...................................................................................... 14
Figure 7. SEM images of core-shell nanofibers fabricated with varying experimental
conditions and pore diameter of control and CAP-treated samples. ......................... 27
Figure 8. Release profiles of “NON-CAP” (FITC-dextran loaded core-shell nanofibers)
and “CAP” (CAP treated FITC-dextran loaded core-shell nanofibers) (N = 1). ...... 28
Figure 9. Water contact angles on CAP-untreated and CAP-treated samples (N = 3), *p <
0.05, **p < 0.01, ***p < 0.001. ................................................................................ 29
Figure 10. Fibroblast (A) and osteoblast (B) proliferation on control samples, Control/CAP
(CAP-treated control samples), BSA (BSA-loaded samples), and BSA/CAP (CAP-
treated, BSA-loaded samples) after 1, 3 and 5 days of culture (N = 9); *p < 0.05, **p
< 0.01, ***p < 0.001. ................................................................................................ 31
Figure 11. Confocal microscopy images of osteoblasts seeded on control samples,
Control/CAP (CAP treated control samples), BSA (BSA loaded samples), and
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BSA/CAP (CAP treated BSA loaded samples) after 1 and 3 days culture. F-actin was
stained red and nuclei was stained green. Scale bar is 50 µm. ................................. 32
Figure 12. ALP activity (A), total protein content (B), calcium content (C), ALP activity
that normalized to total protein (D), and calcium content that normalized to total
protein (E) of MSCs after 7, 14 and 21 days osteogenic differentiation on control
samples, Control/CAP (CAP-treated samples), BSA (BSA-loaded samples), and
BSA/CAP (CAP-treated, BSA-loaded samples) (N = 3), *p < 0.05 ........................ 35
Figure 13. Protein adsorption of control samples and CAP-treated samples (N = 3); *p <
0.05............................................................................................................................ 36
Figure 14. Swelling ratio of control samples, Control/CAP (CAP-treated control samples),
BSA (BSA-loaded samples) and BSA/CAP (CAP-treated, BSA-loaded samples) (N
= 3); *p < 0.05. .......................................................................................................... 36
Figure 15. Pore area distribution of control, samples with 30s CAP treatment, samples with
60s CAP treatment, and samples with 90s CAP treatment. ...................................... 54
Figure 16. Schematic illustration of the fabrication of BMP-2-loaded BSA nanoparticles
stabilized with chitosan (A) and electrospinning of nanoparticle-embedded Poly(ε-
caprolactone)-poly(ethylene glycol) (PCE) copolymer nanofibers (B) (adopted
from57). ...................................................................................................................... 55
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1. Introduction
1.1. Motivation and Background
Tissue deficiency, which refers to the loss or failure of an organ or tissue, is a prevalent
and costly human health care problem 1. Bone tissue deficiencies including bone injuries
or defects are significant clinical problems in tissue deficiencies. Annually, more than half
a million patients receive treatment for bone defects. The expense is more than $2.5 billion
2. The current standard method is to use metallic implants to align bone. Although those
implants are appealing for their mechanical strength, they show numerous side effects like
infection, stiffness, chronic pain and others 3. Autografts are ideal transplants to treat bone
tissue deficiencies since they can reduce the risk of immune rejection. However, the limited
supply of autografts and associated donor site morbidity that result from these
transplantations are significant drawbacks 3.
In 1993, Langer and Vacanti introduced tissue engineering, which is an
interdisciplinary field of engineering involving developed biological substitutes to
maintain, restore, or improve tissue function 1. The bioactive scaffold is one of the
strategies in tissue engineering for creating new tissue. It consists of natural or synthetic
materials to overcome common drawbacks in traditional tissue transplantations. After
culturing cells in the scaffolds, these systems can be implanted or serve as extracorporeal
devices. The application of such scaffolds has demonstrated successful tissue regeneration
throughout the body, including nervous system, cornea, skin, liver, cartilage, bone, muscle,
blood vessel tissue and so on 1, 4.
A major advantage of fibrous bioactive scaffolds that promotes their use in tissue
regeneration is that they effectively mimic the extracellular matrix (ECM). The ECM, a
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fibrous nano- and microscale three-dimensional environment in tissues, is critical to proper
cell and tissue function because it provides both an environment that supports cell growth
and biological cues that direct cellular behavior synchronously 5-6.
1.2. Objectives and Scope
The main objective of this study was to develop core-shell nanofibrous scaffolds that
can not only promote cell proliferation, but can also deliver bioactive proteins in a
controlled manner. The specific aims of this thesis include:
1. To optimize the electrospinning parameters for fabricating uniform and bead-free
polyvinyl alcohol (PVA)/poly (L-lactic acid) (PLLA) core-shell nanofibers.
2. To apply cold atmospheric plasma (CAP) to modify electrospun scaffolds and
observe the nanofiber morphology by using scanning electron microscopy.
3. To measure the water contact angle of the electrospun scaffolds and to analyze the
scaffold’s swelling and water content.
4. To evaluate the scaffold’s protein adsorption by using a bicinchoninic acid protein
assay kit.
5. To load bovine serum albumin (BSA; as a model protein) into the core-shell
nanofiber and to test fibroblasts and osteoblasts proliferation on the electrospun
scaffolds.
6. To observe osteoblast morphology on BSA-loaded core-shell nanofibers by using
confocal microscopy.
7. To conduct osteogenic differentiation of mesenchymal stem cells (MSCs) on the
BSA-loaded nanofibrous scaffold and evaluate the osteoinductivity of the BSA-
3
loaded nanofibrous scaffold by measuring the alkaline phosphatase (ALP) activity,
total protein content and calcium deposition.
8. To load fluorescein isothiocyanate-dextran (FITC-dextran) into the core-shell
nanofibers and to conduct drug release assays to obtain release profiles.
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2. Critical Literature Review
2.1. Bone Grafts and Bone Tissue Engineering (BTE)
Bone, which comprises bone tissue and bone marrow, is an organ with an inherent
capacity to repair and regenerate by itself 7. However, for some severe bone defects caused
by trauma, pathological factures, or compromised blood supply (osteonecrosis), bone fails
to effectively self-heal. These bone tissue deficiencies can result in permanent defects.
Therefore, bone tissue is evidenced to be one of the most frequently transplanted tissues 7.
The major treatment for bone defects remains bone grafting. A bone graft (table 1) is
an implantable material which can promote bone regeneration due to its related biological
functions like osteoinduction, osteoconduction and osteogenesis. The bone grafts can be
autografts, allografts, xenografts or biomaterial substitutes 7.
Table 1. Advantages and disadvantages of three types of bone grafts (modified from 7).
Advantages Disadvantages
Autograft • Osteogenic
• Osteoconductive
• Osteoinductive
• Increased patient morbidity
• Lack of vascularization
• Limited quantity and
availability
Allograft or
Xenograft
• Osteoconductive
• Osteoinductive
• High availability
• No donor site morbidity
• Lacking osteogenicity and
vascularization
• Higher rejection risk
• Risk of disease transmission
• Higher cost
5
Engineered
substitutes
• Ability to improve
osteogenicity and graft
incorporation
• Can use 3D printing
techniques to match the
defect site
• No donor site morbidity
• Osteogenicity is limited by
material porosity
• Variable biodegradability of
different materials
• Poor neovascularization
• Unknown immune response
• Limited mechanical
properties
Both autografts and allografts have significant drawbacks such as morbidity and
infection at associate donor sites and risks of disease transmission (Table 1). Since 1993,
Langer and Vacanti introduced tissue engineering. Engineered substitutes, which combine
a scaffold, cells, and bioactive agents like growth factors have been developed to serve as
an alternative strategy to repair bone deficiency 1. It has given promising results in
providing advanced clinical treatment. The application of a tissue engineered implant can
decrease the numbers of surgeries associated with the removal of implanted allografts and
graft harvesting from other tissues. Therefore, the recovery time and costs could be reduced
due to fewer surgery and risks associated with the surgeries can be decreased. Moreover,
the developed biomimetic scaffolds can contribute to a seamless integration between
existing tissue and engineered implant 3. Figure 1 illustrates the paradigm of bone tissue
engineering.
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Figure 1. Paradigm of Bone Tissue Engineering (adopted from 8).
2.2. The Extracellular Matrix (ECM) of Bone
One major approach in bone tissue engineering is to mimic the extracellular matrix
(ECM), which is a non-cellular macro-molecular network (Figure 2). It not only provides
sites to support the surrounding cells but also provides biological cues that are essential to
cell behaviors such as cell attachment, proliferation, and differentiation 9. The ECM in bone
consists of collagen I fibers and nano-scale hydroxyapatite (HA) 10. The Collagen I has
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both structural and cell signaling functions. It can interact with cell surface receptors such
as integrins, which regulate cells adhesion to the ECM, and also directs mesenchymal
stem cells (MSCs) to differentiate along the osteoblastic lineage 10. Other organic phases
of ECM are composed of hundreds of different proteins such as proteoglycans 10. They also
have essential functions such as the organization and homeostasis of bone. These bone
matrix proteins contribute to the fibrillogenesis of collagen, mineralization, growth factor
release and cell signaling 11.
Figure 2. Schematic of the extracellular matrix (adopted from 12).
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2.3. Nano-fibrous Scaffolds in Bone Tissue Engineering
The goal of the nano-fibrous scaffolds is to create a 3D environment that mimics the
ECM for cells to adhere, proliferate, migrate and differentiate 1. These tissue-engineered
scaffolds require some crucial characteristics, such as high porosity which enables efficient
transportation of different kinds of substances like proteins and waste products. Moreover,
the scaffolds must have a sound mechanical property to withstand the stresses during the
process of tissue neogenesis. Biodegradability and biocompatibility are two other
important characteristics for an ideal scaffold in bone tissue engineering 3.
There are two main methods to fabricate nanofibrous scaffolds: electrospinning (Fig.
3c) and phase separation (Fig. 3a). Other methods like molecular self-assembly (Fig. 3b),
electrohydrodynamic printing, and bacteria-derived hydrogels can also be used to produce
nanofibrous scaffolds. However, they are not widely utilized in bone tissue engineering 3.
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Figure 3. Schematic of current fabrication techniques of nano-fibrous scaffolds. Scale bar: 10 µm (a, c), 500 nm (b) (adopted from 12).
2.4. Electrospinning
Electrospinning, which can produce micro- or nano-fibers is a promising method to
fabricate various bioactive polymeric scaffolds because of its simplicity and versatility 13.
The basic set up of the electrospinning device contains a spinneret, collector and high
voltage source. Due to high surface to volume ratios, these electrospun scaffolds provide
adequate space to promote scaffold-cell interactions 14. In 2005, coaxial electrospinning,
which was first introduced by Sun et al., was a significant improvement of the
electrospinning technique 15. The difference between conventional electrospinning and
coaxial eletrospinning (CE) is that in CE, the spinneret consists of two concentric needles.
10
Two different polymer solutions are injected into the needles separately and then meet at
the tip of the needles. The disadvantage of conventional electrospinning is that it cannot
fabricate nanofibers with non-electrospinnable materials. However, CE can overcome this
drawback by using an electrospinnable material to form the shell 16. Also, the CE can be
utilized to encapsulate drugs and biological agents, such as enzymes and growth factors,
into core polymers for targeted drug delivery 17-18. While water-soluble agents such as
DNA and proteins usually have low solubility in organic solvents, they can be incorporated
into scaffolds (i.e., become electrospinnable) by blending them into water-soluble
polymers. The shell can subsequently control the release kinetics of these materials to
provide optimal conditions for cell growth. Additionally, the shell can also prevent the
loaded substance from an initial burst release 19-21. The materials that are ideal for the CE
will be discussed below.
Figure 4. Schematic diagram of the coaxial electrospinning (modified from 16).
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2.5. Poly (L-lactic acid) (PLLA)
PLLA is a USA Food and Drug Administration (FDA) approved bioabsorbable
polymer for biomedical use. It is a slow crystallizing, semi-crystalline polymer. Its
crystallinity is from 40% to 50%. Its glass transition and melting temperatures are 55-80℃
and 170-180℃ 22. In vivo studies have shown that it can take 2 to 5 years for high molecular
weight PLLA to completely biodegrade 23. It is biodegradable and has robust mechanical
properties 24, thus it is widely used as a material for electrospinning in bone tissue
engineering applications. PLLA can be blended with other bioactive materials such as
gelatin, collagen, or growth factors to increase electropsinnability, or it can be used as the
shell material in CE. Liu et al. blended PLLA with a pentamer-graft-gelatin because PLLA
can act as a support material for scaffolds 25. In order to mimic the ECM of bone, Molamma
et al. blended PLLA with collagen and nanohydroxyapatite. Cell adherence, proliferation
and osteoconduction were improved 26. Schofer et al. developed BMP-2-loaded PLLA
electrospun nanofibers for in vivo bone regeneration 27. The combination of BMP-2 and
PLLA significantly improved in vivo bone regeneration in a rat critical sized calvarial
defect model. However, the release kinetics of BMP-2 were not reported in their paper.
Additionally, PLLA can also be used to form core-shell nanofibers by coaxial
electrospinning. Xu et al. fabricated poly(glycerol sebacate) (PGS)/PLLA core-shell
nanofibrous scaffolds, which demonstrated mechanical properties like soft tissue. In vivo
studies have shown that the poly(glycerol sebacate) (PGS)/PLLA core/shell fibrous
scaffolds can be applied to deliver enteric neural crest (ENC) progenitor cells into the gut
environment 28. Ji et al. developed chitosan/PLLA core-shell nanofibrous scaffolds as a
potential drug carrier in tissue engineering 29. The major disadvantage of PLLA in BTE is
12
that it is very hydrophobic. Thus, it can be difficult for culture media to permeate the
scaffold. Additionally, the absence of sufficient adhesive sites for cell receptors results
in a negative impact on cell attachment 30. Plasma treatments have been shown to improve
the hydrophilicity and cell attachment of polymers 31-34. The details of plasma modification
on PLLA will be discussed below.
2.6. Polyvinyl Alcohol (PVA)
PVA, a semicrystalline hydrophilic polymer with biodegradable and biocompatible
properties 35. It is electrospinnable and water soluble, thus it has already been widely
applied for loading proteins such as gelatin, sericin and bovine serum albumin (BSA) as
well as other substances like antibiotics, silk and chitosan 36-40. However, it has poor
stability and insufficient consistency in aqueous medium 41. CE can be an approach to
overcome this disadvantage by choosing a water insoluble material as the shell material
like PLLA, which has been discussed above.
2.7. Cold Atmospheric Plasma (CAP)
CAP (Fig. 5), which is a cluster of ionized gases where the ion temperature is close to
room temperature, can be used as an effective plasma surface modification (PSM) tool.
CAP contains electrons, reactive oxygen species (ROS), reactive nitrogen species (RNS),
radicals, and UV light 34, 42-43. A previous study demonstrated that the reactive oxygen
species include peroxides, superoxide, hydroxyl radicals, and atomic oxygen. Under CAP
treatment, the –CH3 functional group can react with the charged species, and be converted
to –CH2OH, –CHO, and –COOH (Fig. 6). This oxidization process can change the surface
chemical groups on a PLA scaffold, resulting in the change of surface hydrophilicity and
roughness. A change of nanofiber hydrophilicity and surface morphology significantly
13
enhanced osteoblast and MSC adhesion as well as other functions 33. For example, it was
reported that the attachment of mouse 3T3 fibroblasts on PLLA films was significantly
enhanced 44 and binding sites for mesenchymal stem cell (MSC) integrin receptors
increased after plasma treatments 33, 45.
Figure 5. Schematic of the cold atmospheric plasma setup (adopted form 46, 47).
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Figure 6. Possible mechanisms of chemical and roughness changes caused by CAP on PLA scaffolds (adopted form 33).
2.8. Summary
In summary, the main objective of this study was to develop CAP-modified BSA-
loaded PVA/PLLA core-shell nanofibrous scaffolds. BSA can be used as a model bioactive
protein to be loaded into the core polymer for cell studies, because this protein is
biocompatible, nontoxic, biodegradable, and not immunogenic 48. The hypothesis is that
the CAP treatment can not only promote cell adhesion and proliferation, but also can
impact release the kinetics of loaded substances.
3. Materials and Methods
3.1. Optimization of Coaxial Electrospinning Parameters
PLA
15
Core-shell nanofibers were fabricated by a coaxial electrospinning process (Fig. 4).
This whole process was conducted in the a Doublespinner (Inovenso, Turkey). Poly(L-
Lactic acid) (PLLA, molecular weight 100,000-125,000 Da, Akina, US) was dissolved in
a dichloromethane (DCM, Sigma-Aldrich, USA) and a N,N-Dimethylformamide (DMF,
Sigma-Aldrich, US) mixed solvent. The polymer solution was used to form the core of the
nanofibers and was created by dissolving poly(vinyl alcohol) (PVA, molecular weight
89,000-98,000, Sigma-Aldrich, US) in DI water.
The surface morphology of the nanofibers can be controlled by altering experimental
parameters such as solution concentration, the distance between nozzles and the collector,
the flow rate of the two solutions, and the strength of the applied electric field (applied
voltage) 49. Optimization of these parameters was essential to achieve the optimal
nanofibers with the desired diameter, shape and porosity. This process was conducted by
individually changing the variables mentioned above while keeping all others constant. In
total, 10 experimental runs were performed to predict the optimized conditions (Table 3).
After optimization, bovine serum albumin (BSA, Sigma-Aldrich, US) loaded core-shell
nanofibers were fabricated by adding BSA (0.5%) into a 6% (w/v) PVA in DI water
solution by coaxial electrospinning with the optimized parameters.
16
Table 2. Optimization of coaxial electrospinning parameters.
Experiment number
PLLA (w/v)
PVA (w/v)
Flow rate (mL/hr) Distance (cm)
Voltage (kV)
PLLA PVA
1 10% 6% 3 0.5 15.5 10.5
2 8% 6% 3 0.5 15.5 10.5
3 12% 6% 3 0.5 15.5 10.5
4 10% 6% 1.5 0.5 15.5 10.5
5 10% 6% 2.5 0.5 15.5 10.5
6 10% 6% 3.5 0.5 15.5 10.5
7 10% 6% 3 0.5 13.5 10.5
8 10% 6% 3 0.5 17.5 10.5
9 10% 6% 3 0.5 15.5 8.5
10 10% 6% 3 0.5 15.5 12.5
3.2. CAP Treatment
CAP treatment was conducted to change the nanofibers’ surface morphology, which
can be used to control the drug release profile. Electrospun scaffolds were treated
downstream of a CAP system (Micro-propulsion and Nanotechnology Laboratory (MpNL),
The George Washington University), which was supplied by compressed helium gas. Fig.
5 illustrates the setup of the CAP system. In order to compare the pore size after different
durations of CAP treatment, samples were treated for 30s, 60s, and 90s with a 12 mm
distance between plasma jet and sample surface. Untreated samples were considered as
controls.
17
3.3. Evaluation of Surface Morphology
To determine the surface morphology of the scaffolds, all electrospun scaffolds either
fabricated through the optimization process or CAP-treated were sputter-coated with
platinum. The thickness of the coating was limited to 5 nm. Samples were subsequently
imaged using a scanning electron microscope (SEM) (Hitachi S4800) at a 3 kV accelerating
voltage. After obtaining the SEM images, ImageJ was used to analyze the the pore size.
For each sample, 100 pores on the nanofibers were randonly selected and measured by
ImageJ.
3.4. Measurements of Surface Wettability
The wettability of the nanofiber surface was determined by using a contact angle
analyzer (Data Physics Corp., San Jose, CA, US). CAP treated and non-treated samples
were cut into 1 cm2 square mats. 30 µl droplets of deionized (DI) water were deposited in
random locations on the surface of three 1 cm2 square mats. At different time intervals, the
water contact angle was measured. The surface energy can be calculated from the mean
value of contact angle by using Young–Dupré equation 50-51:
−Δ𝐺 = 𝛾 (1 + cos 𝜃)
ΔG is the free energy of adhesion, γ is the surface tension of water (γ = 0.0720 N/m 52),
and θ is the water contact angle.
3.5. In vitro Drug Release
70 kDa fluorescein isothiocyanate-dextran (FITC-dextran) (Sigma, US), which has a
similar molecular weight to BSA, was loaded into the core fibers by adding it into the PVA
solution (6%, w/v). Ten uniform, 10 mm-diameter mats of each FITC-dextran loaded
electrospun scaffold, as well as 10 CAP-treated (60s) samples, were immersed in 10 mL of
18
DI water in 20 mL glass vials at 37℃ under shaking (120 rpm) for 15 days. 100 µL
solutions of each sample with the released FITC-dextran were then pipetted into 96-well,
clear bottom black microplates (Corning, US) at various time intervals. The concentration
of each released FITC-dextran was evaluated by measuring fluorescence using a
UV/fluorescence spectroscopy (SpectraMax M3 microplate reader, Molecular Devices, US)
with an excitation max of 485 nm and emission max of 525 nm. The addition of fresh DI
water was conducted to replace the volume of withdrawn solutions.
3.6. Osteoblast and Fibroblast Response on Nanofibrous Scaffolds
3.6.1. Cell Culture
Human dermal fibroblasts (HDF) (ATCC CCL-110, US) were cultured in DMEM
(Corning, US) supplemented with 10% fetal bovine serum (ATCC, US) and 1%
penicillin/streptomycin (P/S) (Thermo Fisher, US) in a 37℃ and humid air environment
with 5% CO2. Human osteoblast cells (OB) (Promocell C-12720, Germany) were cultured
in Osteoblast basal medium (Promocell, Germany) supplemented with osteoblast growth
medium supplementMix (Promocell, Germany) and 1% P/S (v/v) in a 37℃ and humid air
environment with 5% CO2. Cell culture medium was changed every two days.
3.6.2. Cell Proliferation Study
Non-loaded samples and BSA-loaded samples were cut into 10 mm-diameter round
mats. After 60s of CAP treatment, four groups including a control, control/CAP (CAP
treated samples), BSA (BSA loaded samples) and BSA/CAP (CAP treated and BSA loaded
samples), were placed in a 24-well plate and sterilized by UV light for 20 min. Each group
was in triplicate. Cell proliferation assays were conducted for both HDF and OB cells under
the same condition. Briefly, scaffolds were soaked in culture medium for 20 min before
19
cell seeding. After removing the soaking medium, cells were seeded onto the samples at a
density of 25,000 cells/mL. After culturing for 1, 3 or 5 days, samples were moved to a
new plate. A 500 µL solution of fresh media and MTS dye (Promega, Madison, WI) at a
ratio of 5:1 was added to each sample. Then, samples containing the MTS solution were
incubated for 3 h at 37°C. Then, 120 µL of the MTS-media solution from each well was
pipetted to a 96-well plate in duplicate. The absorbance of the solution was measured using
a SpectraMax M3 microplate reader at a 490 nm wavelength. The number of cells that
adhered to each sample was obtained by comparing the absorbance values to a standard
curve.
3.6.3. Osteoblast Vinculin Staining
Osteoblasts were seeded on the four sample groups (control, control/CAP, BSA, and
BSA/CAP) in a 24-well plate at a concentration of 50,000 cells/mL. After 1, 3 and 5 days
of culture, samples were washed with PBS to remove non-adhered cells, fixed with 4%
paraformaldehyde for 20 min at room temperature, and washed with PBS twice. Then, 0.1%
Triton X-100 (ThermoFisher, US) in PBS was used to permeabilize the cells and samples
were subsequently washed twice with PBS. Non-specific antigens were then blocked by
incubating samples in a 5% BSA in PBS solution at 4℃ overnight. After washing the cells
twice with PBS, FITC-anti-vinculin (Sigma-Aldrich, US) was diluted (1:800) in a BSA
solution (1% in PBS) and then added to samples for 20 min. Next, Alexa Fluor 568
Phalloidin (ThermoFisher, US) was diluted in a BSA solution (1% in PBS) at a ratio of
1:25. Then, it was applied to cells for 20 min and then washed with PBS twice. Finally, to
stain the nuclei of cells, a solution of 1 mg/mL Hoechst 33258 (ThermoFisher, US) in DI
water was added to samples for 10 min, followed by rising with PBS. Before observing
20
samples with an inverted confocal laser scanning microscopy system (Zeiss 700, US),
samples were mounted on cover glasses using a mounting medium.
3.7. Mesenchymal Stem Cell (MSC) Response on Nanofibrous Scaffolds
3.7.1. MSC Culture
MSCs (ATCC PCS-500-012, US) were cultured in mesenchymal stem cell basal
medium (ATCC, US) supplemented with a mesenchymal stem cell growth kit (ATCC, US)
and 1% P/S (v/v) in a 37℃ and humid air environment with 5% CO2. Cell culture medium
was changed every two days.
3.7.2. MSC Differentiation
MSCs were seeded on four sample groups (control, control/CAP, BSA, and BSA/CAP)
in a 24-well plate at a concentration of 39,000 cells/mL. After culturing for 7 days, old
medium was removed and osteocyte differentiation tool (ATCC, US) with 0.1% P/S (v/v)
was added to provide the medium for differentiation. Osteogenic differentiation medium
was changed twice every week. After 1, 2 and 3 weeks, alkaline phosphatase (ALP) activity,
calcium content and total protein production were tested to evaluate the osteoinductivity of
the nanofibrous scaffolds, as described below.
3.7.3. ALP Activity Measurement
After 1, 2 and 3 weeks differentiation, the cells on the scaffolds were lysed with 0.2%
Triton X-100 (Sigma-Aldrich, US) in DI water by shaking for 20 min at room temperature.
ALP assay kit (QuantiChrom™, BioAssay Systems, US) was used to evaluate the ALP
activity. Reagents were prepared according to manufacturer’s instructions. 50 µL cell
lysates were transferred into a 90 well plate followed by adding 150 µL reagents into
sample wells. Additionally, 200 µL of DI water and 200 µL of Calibrator were added into
21
the 96 well plate. The absorbance was measured on a SpectraMax M3 microplate reader at
405nm at 0 min (ODt=0) and 4 min (ODt=4). The ALP activity was calculated from the
equation which was provided by the manufacturer of ALP assay kit:
ALP activity (U/L)=ODt=4 - ODt=0
ODCalibrator – ODDI water×35.3
3.7.4. Total Protein Content Measurement
The total protein content was measured by using a BCA protein assay kit. The reagents
were prepared according to the manufacturer’s instructions. 25 µL of cell lysates (prepared
as described above) were transferred into a 96 well plate followed by adding 200 µL
reagent into the sample wells. After incubating at 37℃ for 30 min, the absorbance was
measured at 562nm on a SpectraMax M3 microplate reader. The absorbance was converted
to protein content by using an albumin (BSA) standard curve.
3.7.5. Calcium Content Measurement
The calcium content was determined by using a Calcium Assay Kit (QuantiChrom™,
BioAssay Systems, US). 0.5 mL of cell lysates (prepared as described above) were mixed
with 0.5 mL HCl (1.2 M) by shaking at 120 rpm at room temperature overnight to dissolve
the anticipated calcium deposition. Working solutions were prepared according to
manufacturer’s instructions. Then, 5 µL of supernatants were transferred into a 96 well
plate followed by adding 200 µL of working solution into sample wells. After 3 min of
incubation, the absorbance was measured by using a SpectraMax M3 microplate reader.
The absorbance was converted to Ca2+ concentration by using a standard curve.
3.8. Protein Adsorption
Protein adsorption on electrospun scaffolds was measured as in previous studies 53-54.
Control samples and CAP treated samples were punched into 10-millimeter-diameter mats
22
and placed into a 24-well plate. After adding PBS into each well, samples were incubated
at 37℃ overnight. Then, 200 µL of a BSA solution (2 mg/mL in PBS) was added to each
sample after removing PBS. After incubating at 37℃ overnight, PBS was used to wash
samples to remove non-adhered proteins. Finally, the working reagents prepared from the
Bicinchoninic acid (BCA) protein assay kit (Pierce Protein Assay Kit, Thermo Scientific,
US) were added into each sample. After incubating at 37℃for 30 min, 200 µL of the
incubated solution was added into a 96 well clear bottom black microplate (Corning, US).
The absorbance was measured at 562nm on a plate reader. The concentration of adsorbed
BSA was calculated by comparing the absorbance to a standard curve.
3.9. Analysis of Scaffold Swelling and Water Content
Scaffold water content and swelling were investigated as previously described 55.
Briefly, four sample groups (control, control/CAP, BSA, and BSA/CAP) were cut into 10
mm-diameter round mats and were placed in 24-well plates. Samples were incubated at 37℃
in DI water until equilibrium was reached. After removing the water on the surface with
filter paper, wet samples were weighed using a scale (XS105, Mettler Toledo). The fully
swollen scaffold weights (Ww) were recorded for later calculation. Afterwards, wet
samples were dried at 45℃ for 6 hours. The dried scaffold weights (Wd) were then
measured. The swelling ratio and the water content were calculated from the equations:
Swelling ratio % =(𝑊w−𝑊d)
𝑊d×100
Water content % =(𝑊w−𝑊d)
𝑊w×100
23
3.10. Statistical Analysis
All experiments were performed in triplicate, unless otherwise specified. Data are
expressed as mean ± standard deviation. Statistical analysis was carried out using student
t-tests with statistical significance considered at p < 0.05.
4. Results
4.1. Surface Morphology of Nanofibers
SEM images illustrate the surface morphology of fabricated nanofibers with varying
experimental conditions (Fig. 7). When a concentration of cell solution of 8% (w/v) PLLA
was used, the morphology of the nanofibers was non-uniform. Also, bead defects on
nanofibers could be observed. (Fig. 7B). No significant change in the nanofiber
morphology was observed when the concentration of the shell solution was increased from
10 % (w/v) PLLA to 12% (w/v) PLLA (Fig. 7C). Moreover, a large disparity in fiber
diameter was observed when the flow rate of the shell solution decreased below 3.0 mL/h
(1.5 and 2.5 mL/h) (Fig. 7D and E). As the shell solution flow rate increased to 3.5 mL/h,
the nanofiber diameters remained disparate (Fig. 7F). In addition, nanofiber morphology
was markedly influenced by the distance between the nozzle and the collector. At a distance
of 13.5 cm, fibers occasionally adhered to each other and the nanofiber morphology was
uneven (Fig. 7G). However, at an increased distance of 17.5 cm, the fabricated nanofibers
were similarly non-uniform, indicative of an ideal distance between 13.5 and 17.5 (Fig.
7H). Beaded fibers were observed with a voltage as low as 8.5 kV (Fig. 7I). Nanofibers
were also asymmetrical if the voltage increased to 12.5 kV, again indicating an ideal
voltage between 8.5-12.5 kV (Fig. 7J). Ideal nanofibers that were both straight and having
uniform diameters were obtained using a shell polymer concentration of 10% (w/v) PLLA
24
with a 3 mL/h flow rate, a core polymer concentration of 6% (w/v) PVA with a 0.5 mL/h
flow rate, a distance between the nozzle and collector of 15.5 cm, and a voltage of 10.5
kV (Fig. 7A).
Fig. 7K, L shows SEM images of the nanofibers both before (Fig. 7K) and after CAP
treatment (Fig. 7L). These images demonstrated that surface morphology of the shell was
significantly changed after CAP modification. Additionally, nanofiber pore size was
enlarged after CAP treatment (Fig. 7L) and pores adapted a cave-like structure that was
more tridimensional (Fig. 7L). In addition, some parts of the shell peeled off after CAP
treatment (Fig. 7L). Although pores were observed on untreated nanofiber surfaces, the
surface morphology of these pore was smoother than the CAP treated pores (Fig. 7K).
ImageJ was used to analyze the pore size. The change of pore size after CAP treatment
can be found in Figure 7M. Each value is the mean value of pore size distributions for
different samples (Fig. 15 in appendix). The pore size was increased after CAP treatment.
The pore diameter of the control was about 125 nm. It increased to 175 nm after 30s CAP
treatment, and 225 nm after 60s CAP treatment. However, there was no significant increase
in pore size after 90s CAP treatment compared with the pore size of the 60s CAP-treated
samples.
25
Table 3. Morphology of nanofibers fabricated by coaxial electrospinning with different parameters.
Experiment number
PLLA (w/v)
PVA (w/v)
Flow rate (mL/hr)
Distance (cm)
Voltage (kV)
Fiber Morphology
PLLA PVA
A 10% 6% 3 0.5 15.5 10.5 Optimized
B 8% 6% 3 0.5 15.5 10.5 Bead defects
C 12% 6% 3 0.5 15.5 10.5 Uniform
D 10% 6% 1.5 0.5 15.5 10.5 Not uniform
E 10% 6% 2.5 0.5 15.5 10.5 Not uniform
F 10% 6% 3.5 0.5 15.5 10.5 Not uniform
G 10% 6% 3 0.5 13.5 10.5 Not uniform
H 10% 6% 3 0.5 17.5 10.5 Not uniform
I 10% 6% 3 0.5 15.5 8.5 Bead defects
J 10% 6% 3 0.5 15.5 12.5 Not uniform
26
Figure 7. SEM images of core-shell nanofibers fabricated with varying experimental conditions (Table 3.). A. Optimized parameters: a shell polymer concentration of 10% (w/v) PLLA with a 3 mL/h flow rate, a core polymer concentration of 6% (w/v) PVA with a 0.5 mL/h flow rate, a distance between the nozzle and collector of 15.5 cm and a voltage of 10.5 kV, B. low concentration of PLLA (8%), C. high concentration of PLLA (10%), D. low PLLA flow rate (1.5 mL/hr), E. low PLLA flow rate (2.5 mL/hr), F. high PLLA flow rate (3.5 mL/hr), G. short nozzle-to-collector distance (13.5 cm), H. long nozzle-to-collector distance (17.5 cm), I. low voltage (8.5 kV), J. high voltage (12.5 kV), K. the optimized sample (A) without CAP treatment (higher magnification), L. the optimized sample (A) with CAP treatment (higher magnification).
27
Figure 7M. Pore diameter of control and CAP-treated samples (each value was calculated from mean value of pore area distribution in appendix).
4.2. In vitro Drug Release
To determine the drug release capability of the fabricated nanofiber scaffolds, the
release of FITC-dextran loaded core-shell nanofibers was studied and its release profile is
shown in Fig. 8. Scaffolds without CAP treatment released approximately 20% of the
loaded FITC-dextran during the initial release stage (< 2 hours). A maximum release was
observed at the 48-hour time point. The total FITC-dextran release over a 193 hour time
period was about 45%. Alternatively, CAP treated samples were found to release over 50%
of the loaded FITC-dextran during the initial release stage (< 2 hours). Moreover, it
sustained this release for 120 hours when it reached its maximum drug release. The total
release was 100%. Drug profile plots demonstrate that the initial dextran release speed was
significantly accelerated for the CAP treated samples compared to non-CAP-treated
samples (Fig. 8). Specifically, CAP-treated samples released almost 30% more loaded
drugs than untreated samples at the 2-hour time point. Additionally, the maximum drug
M
28
release greatly increased after CAP treatment. CAP-treated samples released almost 100%
of the loaded FITC-dextran, more than twice the release from the untreated samples.
Surprisingly, CAP-treated samples continued to release FITC-dextran until 120 hours,
which was almost three times longer than that of the untreated scaffolds.
Figure 8. Release profiles of “NON-CAP” (FITC-dextran loaded core-shell nanofibers) and “CAP” (CAP treated FITC-dextran loaded core-shell nanofibers) (n = 3). Data = mean +/- SEM.
4.3. Water Contact Angle
Contact angles on the CAP-treated and -untreated electrospun scaffolds were measured
to evaluate their wettability. CAP-untreated samples were recorded to have water contact
angles over 100° while CAP-treated samples possessed a contact angle of approximately
50° (Fig. 9). This indicates that scaffold surface hydrophilicity was significantly increased
after CAP treatment. The free energy of adhesion of CAP-modified samples (ΔGCAP = -
3.55 mN/m) is higher than that of the control (ΔGcontrol = -11.80 mN/m).
29
Figure 9. Water contact angles on CAP-untreated and CAP-treated samples (n = 3), *p < 0.05.
4.4. Osteoblast and Fibroblast Response on Nanofibrous Scaffolds
To determine the cytocompatibility of the fabricated scaffolds, the attachment,
proliferation, and morphology of fibroblasts and osteoblasts on four sample groups (control,
control/CAP, BSA, and BSA/CAP) were investigated. The fibroblast and osteoblast
proliferation results after days 1, 3 and 5 are shown in Fig. 10. Fibroblasts (Fig. 10A)
proliferated on control and control/CAP samples at all time points. On day 3 and day 5, an
obvious increase in cell density was observed between BSA samples and control while
proliferation on control/CAP samples was significantly higher compared to BSA samples.
For example, on day 5 the cell density on BSA samples was approximately 24,000
cells/cm2 while the control/CAP samples had 32,000 cells/cm2 on day 5. Furthermore,
CAP/BSA samples always had the highest cell density on each day compared to other
30
groups. For instance, the cell density on CAP/BSA samples was 20,000 and 40,000
cells/cm2 on day 3 and day 5, respectively, which was almost twice the cell density on
control samples.
For osteoblasts (Fig. 10B), BSA samples didn’t significantly promote osteoblast
proliferation. However, CAP/BSA samples displayed significant increases in cell densities
compared to control scaffolds on day 3 and day 5. For example, BSA/CAP samples had
2,500 cell/cm2 more than control on day 3 and 5,000 cell/cm2 more than control on day 5.
Moreover, an enhancement in cell proliferation was observed between control and
control/CAP samples on day 5. Also, control/CAP samples and BSA/CAP samples had an
approximate 2,500 cells/cm2 increase in cell density when compared to control and BSA
samples.
To examine osteoblast morphology, laser scanning confocal microscopy was used. As
shown in Fig. 11, the red area indicates the f-actin, a major cytoskeletal component. The
blue area indicates the nucleus. On day 1, osteoblasts on BSA/CAP samples had
significantly larger coverage of the cytoskeleton compared to other samples. On day 5,
cells on control/CAP and BSA/CAP samples had higher cytoskeletal coverage than CAP-
untreated samples. Specifically, clusters of cells were observed on the BSA/CAP samples.
31
Figure 10. Fibroblast (A) and osteoblast (B) proliferation on control samples, Control/CAP (CAP-treated control samples), BSA (BSA-loaded samples), and BSA/CAP (CAP-treated, BSA-loaded samples) after 1, 3 and 5 days of culture (n = 9); *p < 0.05, **p < 0.01, ***p < 0.001. Data = mean +/- SEM.
Day 1
Day 3
Day 5
0250050007500
100001250015000175002000022500
Cel
l Den
sity
cel
ls/c
m2
ControlControl/CAPBSABSA/CAP
* * ***
B
Day 1
Day 3
Day 5
0
8000
16000
24000
32000
40000
48000
56000
Cel
l Den
sity
cel
ls/c
m2
Control
Control/CAPBSA
BSA/CAP
*** ***
****
******** ***
A
32
Figure 11. Confocal microscopy images of osteoblasts seeded on control samples, Control/CAP (CAP treated control samples), BSA (BSA loaded samples), and BSA/CAP (CAP treated BSA loaded samples) after 1 and 3 days of culture. F-actin was stained red and nucleus were stained blue. Scale bar is 50 µm.
4.5. MSC Response on Nanofibrous Scaffolds
The response of MSCs on the four groups of nanofibrous scaffolds (control,
control/CAP, BSA, and BSA/CAP) during osteogenic differentiation can be reflected by
ALP activity, total protein and calcium content. For ALP activity (Fig. 12A), BSA/CAP
samples always had significantly higher ALP activity than control and control/CAP groups
from day 7. BSA/CAP ALP activity was significantly higher than the BSA group after 14
days and 21 days of osteogenic differentiation. On day 21, the control/CAP, BSA and
BSA/CAP groups had significantly higher ALP activity than the control samples. The total
protein assays showed similar results (Fig. 12B). After 7, 14 and 21 days of osteogenic
differentiation, the total protein content from BSA/CAP was significantly higher than other
groups. Also, control/CAP and BSA groups had significantly higher protein content than
33
control. For the calcium content (Fig. 12C), there was no significant difference between
the four types of samples after 7 days and 14 days of osteogenic differentiation. However,
after 21 days, the BSA/CAP group had significantly higher Ca2+ concentration than control
and control/CAP groups. Also, the control/CAP and BSA groups had significantly higher
calcium content than control.
Fig. 12D and Fig. 12E illustrate the ALP activity and calcium content that were
normalized to the total protein content. These results suggest that, for all the protein
synthesized by osteogenic differentiated MSCs, a larger amount of it was ALP on the
BSA/CAP samples, however, there was a lower percentage calcium per total protein on the
BSA/CAP samples. It could be due to the higher amount of total protein of synthesized by
the differentiated MSCS on BSA/CAP groups. Results of ALP activity and calcium content
assays suggested that CAP modification improved the nanofibrous scaffold
osteoinductivity properties. Especially, the BSA/CAP showed the most significant
osteoinductive results on MSCs differentiation.
Figure 12. ALP activity (A) of MSCs after 7, 14 and 21 days osteogenic differentiation on control samples, Control/CAP (CAP-treated samples), BSA (BSA-loaded samples), and BSA/CAP (CAP-treated, BSA-loaded samples) (n = 3), *p < 0.05. Data = mean +/- SEM.
Day 7
Day 14
Day 21
0
20
40
60
ALP
act
ivity
(IU/L
= µ
mol
/(L·m
in))
ConrolControl/CAPBSABSA/CAP
*** *
****
**
A
34
Figure 12. Total protein content (B), calcium content (C), and ALP activity that normalized to total protein (D) of MSCs after 7, 14 and 21 days osteogenic differentiation on control samples, Control/CAP (CAP-treated samples), BSA (BSA-loaded samples), and BSA/CAP (CAP-treated, BSA-loaded samples) (n = 3), *p < 0.05. Data = mean +/- SEM.
Day 7
Day 14
Day 21
0
50
100
150
200
250
Pro
tein
con
cent
ratio
n (µ
g/m
l)
ControlControl/CAPBSABSA/CAP
***
**
***
**
B
**
Day 7
Day 14
Day 21
0.0
0.5
1.0
1.5
2.0
Cal
cium
con
cent
ratio
n (m
g/m
l)
ControlControl/CAPBSABSA/CAP
****
C
Day 7
Day 14
Day 21
0.0
0.1
0.2
0.3
0.4
0.5
ALP
act
ivity
/ to
tal p
rote
in (µ
g/m
l)
ConrolControl/CAPBSABSA/CAP
** *
**
** ** *
D
35
Figure 13. Calcium content that normalized to total protein (E) of MSCs after 7, 14 and 21 days osteogenic differentiation on control samples, Control/CAP (CAP-treated samples), BSA (BSA-loaded samples), and BSA/CAP (CAP-treated, BSA-loaded samples) (n = 3), *p < 0.05. Data = mean +/- SEM.
4.6. Protein Adsorption and Water Swelling
A protein adsorption assay, using BSA as a model protein, was conducted to investigate
the influence of CAP-treatment on protein adsorption on the nanofibrous scaffolds. In order
to avoid the influence of the loaded BSA, only non-BSA loaded samples were tested.
Protein adsorption increased after CAP treatment. As shown in Fig. 13, control samples
displayed protein adsorption lower than 300 µg/mL per sample. However, CAP-treated
samples had approximately 500 µg/mL of protein adsorption per sample. For the water
adsorption study, the swelling ratio of electrospun scaffolds was also significantly
increased after CAP treatment. As shown in Fig. 14, both CAP-treated samples
(control/CAP and BSA/CAP samples) had swelling ratios of approximately 200%.
Day 7
Day 14
Day 21
0.00
0.02
0.04
0.06
Cal
cium
con
cent
ratio
n (m
g) /
tota
l pro
tein
(µg)
ConrolControl/CAPBSABSA/CAP
***
***
E
36
Nevertheless, untreated samples only displayed a swelling ratio of approximately 100%.
No significant difference was observed in water content of all groups of samples.
Figure 14. Protein adsorption on control samples and CAP-treated samples (n = 3); *p < 0.05. Data = mean +/- SEM.
Figure 14. Swelling ratio of control samples, Control/CAP (CAP-treated control samples), BSA (BSA-loaded samples) and BSA/CAP (CAP-treated, BSA-loaded samples) (n = 3); *p < 0.05. Data = mean +/- SEM.
Control
Control/C
AP0
200
400
600
800
Pro
tein
con
cent
ratio
n (µ
g/m
l) *
Swellin
g ratio
Wat
er co
ntent
0
100
200
300
Sw
ellin
g R
atio
Control
Control/CAP
BSA
BSA/CAP
* *
37
5. Discussion
5.1. Optimization of Coaxial Electrospinning Parameters
In this study, optimization of a coaxial electrospinning technique was first performed
to fabricate uniform and bead-free nanofibers. Bead-free nanofibers were specifically
formed as beads can negatively affect stable drug release. The morphology of nanofibers
was modified by adjusting coaxial electrospinning parameters including polymer
concentration, solution flow rate, distance between nozzles and the collector, and applied
voltage 49. According to optimization experiments, the polymer concentration had a
significant impact on the spinnability of the solution. If the polymer concentration was too
low, fibers broke and split into droplets before attaching to the collector due to reduced
surface tension. This was evidenced by the bean-shaped fibers found in Fig. 7B. Moreover,
nanofiber size and shape were also impacted by solution flow rate. If the flow rate was too
slow, the injected solution couldn’t sufficiently replace the solution ejected by the nozzle.
Therefore, the Taylor cone (a cone formed on the tip of the spinneret during electrospinning)
couldn’t be maintained due to the insufficient supply of polymer solution 56. The unstable
Taylor cone resulted in an uneven distribution of fiber diameter, which can be observed in
Fig. 7D, E. The increase in flow rate contributed to the augmentation of fiber size, which
can be found in Fig. 7F. In addition, the distance between the nozzle and collector played
a very important role in nanofiber morphology. If the distance was too long, the traveling
time was not long enough for the solvent to evaporate. Thus, fibers were dissolved by the
non-volatilized solvent when they were collected on the collector. This resulted in the
production of fibers that were attached to each other (Fig. 7G). Also, a long capillary-to-
collector distance could cause the reduction of nanofiber diameter (Fig. 7H). In addition,
38
the applied voltage also affected the morphology of nanofibers. The formation of beaded
fibers (Fig. 7I) was observed in scaffolds that were fabricated with a weak electric field
strength. If the applied voltage was too low, the Taylor cone became a pendent drop at the
tip of the capillary. The abnormal shape of the Taylor cone subsequently contributed to the
formation of bead defects, which can be detrimental to drug release. When the voltage
increased, the fiber diameter decreased (Fig. 7J). However, if the electric field was too
strong, beaded fibers developed because the fiber jet was ejected within the capillary 49.
5.2. Cell Behavior
In order to promote cell adhesion and proliferation, CAP treatment was introduced to
modify the nanofiber surface morphology. As a previous study demonstrated, the collision
of charged particles and accumulated partial heat converted the scaffold’s surface structure
from the mirco-scale to nano-scale, resulting in larger surface area and surface energy 33.
In this study, CAP modification similarly enlarged the pores on the fiber surface. The
increased pore size and the nano-scale structure contributed to a larger specific surface area
compared with untreated surfaces. The higher surface area not only provided for more sites
for cell attachment, but also promoted protein adsorption to the scaffold, which elevated
scaffold protein adsorption (Fig. 13) and water swelling (Fig. 14), which could further
enhance beneficial cellular behavior.
Another scaffold characteristic that resulted from CAP treatment was the increase of
nanofiber surface hydrophilicity as well as the free adhesion energy, which are proved to
play an important role in cell attachment 31. As J. Yang et al. mentioned in their study,
receptors on the cell membrane are of vital importance in cell adhesion to ECM. Cells sense
their surroundings through these receptors. The process of binding between receptors and
39
ECM molecules is a crucial step in cell attachment. CAP treatment can increase the
hydrophilicity and surface energy of a scaffold 33, which in turns alter initial protein
adsorption to favor cell adhesion. As previously illustrated, CAP treatment can change the
chemicals on the surface by an oxidization process. Due to this oxidization, some of the
methyl groups of the polymer can be changed into hydrophilic chemical functional groups
like hydroxyl methyl. These polar groups increase nanofiber surface hydrophilicity. After
protein adsorption, the scaffold provides sufficient binding sites for receptors during cell
adhesion 31, 33. The observed proliferation results (Fig. 10) and cell phenotype observation
images (Fig. 11) confirmed that CAP treatment could improve fibroblast and osteoblast
attachment and proliferation.
Most importantly, CAP treatment not only increased fiber surface hydrophilicity and
nano-structure, but also enhanced drug release, which is another novelty of this study. After
CAP treatment, surface pore size of the nanofiber increased, thus increasing total surface
area. This characteristic resulted in an increased diffusion rate of the drug-loaded core
polymer. As shown by SEM images (Fig. 7K, L) and the drug release profile (Fig. 8), CAP-
modified samples had larger surface pores and could release the proteins faster and more
completely than untreated samples. In another study, Shalumon et al. compared the BMP-
2 release profile of two types of core-shell nanofibers with different shell thickness. They
found that thin shell nanofibers had higher rates of BMP-2 release than thick shell
nanofibers. However, there was no significant promotion of cell proliferation by altering
shell thickness 57. In contract, our study introduced CAP treatment to control the release
kinetics. This method of controlling drug release provides more flexibility and simplicity
than changing shell thickness. Additionally, when drug release was enhanced, more growth
40
factors in the core polymer could be released. Due to the increased concentration of growth
factor around the surface of the scaffolds, cell attachment and proliferation were
significantly promoted (Fig. 10). Moreover, MSCs responses on the nanofibrous scaffold
(Fig. 12) showed that CAP treatment significantly improved the nanofibrous scaffold’s
osteoinductivity. The possible reason is that the CAP modification could also enhance the
physical cues of osteogenic differentiation which may be due to the increased
hydrophilicity and surface roughness.
6. Conclusions
In this study, a core-shell nanofiber was designed for use as a novel drug delivery
system. Uniform and bead-free nanofibers were fabricated by optimizing coaxial
electrospinning parameters. A CAP treatment was introduced to increase scaffold surface
hydrophilicity and nanofiber surface pore size and specific surface area. Moreover, scaffold
protein and water adsorption as well as drug release were significantly enhanced after
treatment. Collectively, these features promoted the adhesion and proliferation of
fibroblasts and osteoblasts. Also, the CAP modification improved the osteoinductivity of
nanofibrous scaffolds. These convincing results suggest that CAP-modified and bioactive
protein-loaded core-shell nanofibers have a great potential for drug delivery studies and
bone tissue engineering applications.
41
7. Future Directions
The promising results of this study suggested that additional research directions
should be considered as the core-shell nanofibers in future studies as described below.
7.1. Bone Tissue Engineering
Hydroxylapatite (HA) can be loaded into PLLA in order to increase the
osteoinductivity of core-shell nanofibers, because it is a vital component in natural bone 58.
BMP-2 can be loaded into PVA to improve the osteoinductivity since this study has proved
that bioactive protein-loaded core-shell nanofibers show promising properties for
osteoblast proliferation and MSC differentiation. For example, scientists have discovered
a novel way to deliver BMP-2 by synthesizing BMP-2 loaded BSA nanoparticles (BNPs)
(156 ± 18 nm). This nanoparticle-embedded electrospun nanofiber can achieve the
controlled delivery of BMP-2 59 (Fig. 16 ). If the BNPs was loaded into the core polymer,
the exposure of the BNPs can be controlled by using CAP treatment which has been proved
that can change the pore size from 125 nm to 225 nm on PLLA.
7.2. Wound Healing
Electrospinning has shown advantages in bafbrication of nanofibrous membranes for
wound healing. It can not only achieve antibacterial properties, but also achieve rapid
hemostasis for wound healing 60.
Antibiotics like Teracycline, Mupirocin, Gentamicin, and Triclosan can be ideal
antibacterial agents to be loaded into PLLA, because many studies have been conducted to
load these antibiotics into polymers like Poly(lactic acid) (PLA) or PLA/Polycaprolactone
(PCL), and PLA/collagen composites. These studies showed that the antibiotics-loaded
42
nanofibers have effective antibiotic activity against bacteria like Staphylococcus aureus,
Escherichia coli, Staphylococcus epidermidis, and Pseudomonas aeruginosa 61-64.
Nanoparticles like silver nanoparticles (AgNPs) and titanium dioxide (TiO2)
nanoparticles (TiO2NPs) have been loaded into PVA or PVA/chitosan composite in
electrospun nanofibers, and have been shown to have promising antibiotic activity against
Escherichia coli, Staphylococcus aureus, and Klebsiella pneumoniae 65-67.
All of the studies discussed above suggested that antibiotics can be loaded into the
PLLA and nanoparticles like AgNPs or TiO2NPs can be loaded into PVA to enhance the
antibacterial properties of electrospun scaffolds.
43
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53
APPENDIX
Figure 15. Pore area distribution of control (A), and samples with 30s CAP treatment (B).
0 5000 10000 15000 20000 25000 30000 350000
5
10
15
20
25
&
&
Distribution
Pore&area&(nm2)
Sample9control
Mean=&12349.474464,&SD=6477.9241692387
0 10000 20000 30000 40000 50000 60000 700000
5
10
15
20
25
30
35
)
)
Distribution
Pore)area)(nm2)
Sample<CAP<30s
Mean=)24045.403463,)SD=13418.635840906
A
B
54
Figure 15. Samples with 60s CAP treatment (C), and samples with 90s CAP treatment (D).
0 20000 40000 60000 80000 100000 120000 140000 160000 1800000
5
10
15
20
25
30
35
40Sample/CAP/60s
4
4
Distribution
Pore4area4(nm2)
Mean=440852.708283838,4SD=28986.818198802
0 50000 100000 150000 2000000
5
10
15
20
25
30
35
40
Sample-CAP-90s
3
3
Distribution
Pore3area3(nm2)
Mean=342042.910137864,3SD=41706.90884866
C
D
55
Figure 16. Schematic of the preparation of BMP-2-loaded BSA nanoparticles (A) and the fabrication of nanoparticle-embedded Poly(ε-caprolactone)-poly(ethylene glycol) copolymer nanofibers by electrospinning (B) (adopted from 59).