Electrospun nanofibrous structure: A novel scaffold for
tissue engineering
Wan-Ju Li,1,4* Cato T. Laurencin,2 Edward J. Caterson,4 Rocky S. Tuan,4,* Frank K. Ko1,3
1
School of Biomedical Engineering, Science and Health Systems, Drexel University, Philadelphia, Pennsylvania 19104
2
Department of Chemical Engineering, Drexel University, Philadelphia, Pennsylvania 19104
3
Department of Materials Engineering, Drexel University, 31st and Market Street, Philadelphia, Pennsylvania 19104
4
Department of Orthopaedic Surgery, Curtis 501, Thomas Jefferson University, Philadelphia, Pennsylvania 19107
Received 26 July 2001; revised 21 November 2001; accepted 5 December 2001
Abstract: The architecture of an engineered tissue substitute plays an important role in modulating tissue growth. A
novel poly(D,L-lactide-co-glycolide) (PLGA) structure with a
unique architecture produced by an electrospinning process
has been developed for tissue-engineering applications.
Electrospinning is a process whereby ultra-fine fibers are
formed in a high-voltage electrostatic field. The electrospun
structure, composed of PLGA fibers ranging from 500 to 800
nm in diameter, features a morphologic similarity to the
extracellular matrix (ECM) of natural tissue, which is characterized by a wide range of pore diameter distribution,
high porosity, and effective mechanical properties. Such a
structure meets the essential design criteria of an ideal en-
gineered scaffold. The favorable cell–matrix interaction
within the cellular construct supports the active biocompatibility of the structure. The electrospun nanofibrous structure is capable of supporting cell attachment and proliferation. Cells seeded on this structure tend to maintain phenotypic shape and guided growth according to nanofiber
orientation. This novel biodegradable scaffold has potential
applications for tissue engineering based upon its unique
architecture, which acts to support and guide cell growth. ©
2002 Wiley Periodicals, Inc. J Biomed Mater Res 60: 613–621,
2002
INTRODUCTION
material interactions. Biologic functioning is regulated
by biologic signals from growth factors, extracellular
matrix (ECM), and the surrounding cells.3 ECM molecules surround cells to provide mechanical support
and regulate cell activities.4
The ultimate goal of the scaffold design is the production of an ideal structure that can replace the natural ECM until host cells can repopulate and resynthesize a new natural matrix. To achieve this goal, the
scaffold material must be selected carefully, and the
scaffold architecture must be designed to insure that
the seeded cells are biocompatible with the engineered
scaffold.
Biocompatibility can be classified into surface and
structural biocompatibility, which are determined by
the selection of the material and the architecture of the
scaffold, respectively.5 Surface biocompatibility is associated with the surface chemistry of the material.
The chemical characteristics of a material surface will
mediate the adsorption of biologic molecules that
regulate cell activities, such as adhesion and migration.6
The surface chemistry of a tissue-engineered scaffold is dependent upon the type of the material, rang-
Tissue engineering for some time has being recognized as a promising alternative to donor tissues,
which often are in short supply. The promise is that
biologic function lost in host tissues will be able to be
restored and maintained by tissue engineering.1 Although various structures of engineered tissue scaffolds have been developed for tissue replacement, the
goal of producing a clinically useful tissue scaffold
still is far from being realized.
An ideal tissue-engineered scaffold should be mechanically stable and capable of functioning biologically in the implant site.2 Mechanical stability is dependent primarily on the selection of the biomaterial,
the architectural design of the scaffold, and the cell–
*Present address: Cartilage Biology and Orthopaedics
Branch, National Institute of Arthritis, and Musculoskeletal
and Skin Diseases, National Institutes of Health, Bethesda,
MD 20892
Correspondence to: F. K. Ko at Department of Materials
Engineering; e-mail: fko@coe.drexel.edu
© 2002 Wiley Periodicals, Inc.
Key words: tissue engineering; electrospinning; PLGA; scaffold; mesenchymal stem cell
614
ing from natural biopolymers to synthetic polymers.
The most commonly used natural biopolymers include demineralized bone matrix, agarose, collagen,
hyaluronan, basement membrane, and alginate.7–12
Synthetic polymers that are used include degradable
polyesters, such as polyglycolic acid (PGA), polylactic
acid (PLA), and their copolymers, poly (D,L-lactide-coglycolide) (PLGA).13 These biodegradable polymers
have a long history of clinical use and currently are
used in various tissue engineering applications.14,15
Structural biocompatibility is affected by the physical morphology of a scaffold, primarily by its architecture and the dimensions of its building components. The dimensions of the building components of
a scaffold are important factors in regulating cell activities.16,17 Cell behavior is known to be regulated by
the physical properties of an engineered scaffold, such
as the architecture and topography. Previous studies
have shown that cell proliferation is influenced by the
architectural scale of the structure and that adhesion is
affected by the topography of the material.18–20
In connective tissue, ECM is composed of two main
classes of macromolecules, ground substances (proteoglycans) and fibrous proteins (collagens),21 that together form a composite-like structure. Collagens embedded in proteoglycans maintain structural and mechanical stability. The collagen fibrous structure is
organized in a three-dimensional fiber network composed of collagen fibers that are formed hierarchically
by nanometer-scale multi-fibrils.22 Therefore, ideally,
the dimensions of the building blocks of a tissueengineered scaffold should be on the same scale with
those of natural ECM.
A novel scaffold, the electrospun nanofibrous structure, is introduced in this study. This structure, produced by an electrospinning process, is a nonwoven,
three-dimensional, porous, and nano-scale fiber-based
matrix. The electrospinning process was first introduced in the early 1930s and has been continuously
investigated to date.23–31 This process is capable of
producing ultra-fine fibers by electrically charging a
suspended droplet of polymer melt or solution.25,26
In this technique, polymer solution is drawn from
the capillary, forming a suspended droplet at the tip of
the capillary by force of gravity or mechanical pumping combined with electrostatic charge. The polymer
jet is initiated when the electrostatic charge overcomes
the surface tension of the droplet. Nanofibers are
formed by the narrowing of the ejected jet stream as it
undergoes increasing surface charge density due to
evaporation of the solvent.32 The nanofibrous structure produced by the electrospinning process has a
high surface area-to-volume ratio, providing more
substrate for cell attachment (and therefore a higher
cell density per unit of space) compared to other structures.
The electrospinning process has been used in vari-
LI ET AL.
ous applications.27,28 To explore the applications of
this structure in tissue engineering, the properties of
the structure, including porosity and mechanical
properties, were characterized and cell activities
within the nano-scale fiber-based scaffolds were investigated. Two cell types—fibroblasts and bonemarrow-derived mesenchymal stem cells (MSCs)—
were applied on the scaffolds in this study. Fibroblasts
are widely distributed in skin and tendons, and bonemarrow-derived mesenchymal stem cells are pluripotent and capable of differentiating into different connective tissue cell types, such as the chondrocyte, osteoblast, adipocyte, and myoblast.33 These two cell
types were used in this study because replacements of
skin and cartilage tissue are the possible candidates
for future applications of this scaffold.
MATERIALS AND METHODS
Electrospun nanofibrous structures
Electrospun nanofibrous structures were fabricated by the
electrospinning process using the apparatus schematically
shown in Figure 1. A polymer solution was prepared by
dissolving 1 g of copolymer poly(D,L-lactide-co-glycolide)
[85:15; PLGA] (Purac, Lincolnshire, IL) in 20 mL of organic
solvent mixture composed of (1:1) tetrahydrofuran (THF;
Fisher, Pittsburgh, PA) and dimethylformamide (DMF;
Sigma, St. Louis, MO) and mixing it well by vortexing the
mixture overnight.
For the process of electrospinning, polymer solution was
placed in a 20-mL glass syringe fitted with an 18-G needle.
The syringe was fixed at the support at a 45-degree angle
down-tilted from horizontal. Eighteen kilovolts were provided by the high voltage power supply (Gamma High Volt-
Figure 1. Scheme of the electrospinning apparatus: (A)
glass syringe containing polymer solution; (B) nanofiber jet;
(C) copper collecting plate; and (D) power supply. See Materials and Methods for additional details.
ELECTROSPUN NANOFIBROUS SCAFFOLD
age Research, Ormond Beach, FL) and employed at a distance of 20 cm between the copper collecting plate (cathode)
and the needle tip (anode). The polymer solution was drawn
from the syringe, forming a pendant drop at the tip of the
needle by the combining force of gravity and electrostatic
charge. A positive-charged jet ejected from the drop splayed
to the negative-charged target. A nanofibrous structure was
formed on the collecting plate and then carefully removed
for subsequent use.
Physical property characterization
Porosity
Pore diameter distribution, total pore volume, total pore
area, and porosity of the structure were measured by the
AutoPore III mercury porosimeter (Micromeritics Instrument Co., Norcross, GA). Sample reparation and procedures
for measurement were conducted following the instructions
provided by the manufacturer. Briefly, the 1-mm thick electrospun structure was cut into 2 × 5-cm rectangular shapes
and weighed. A sample was placed in the cup of the penetrometer (£s/n-14, 3 Bulb, 0.412 Stem, Powder), which was
closed by tightening the cap. The penetrometer, along with
a sample, was sent into the pressure chamber of the porosimeter for measurement of pore properties.
Tensile property
The 1-mm thick nanofibrous structure was cut into 1 ×
6-cm rectangular shapes, and tensile properties were characterized by the Kawabata Evaluation System (KES-G1, Kato
Tech Co., Japan). The method of tensile determination followed standard mechanical testing methods for fabric materials. The 1 × 6-cm × 1-mm specimen was vertically
mounted on two 1 × 1-cm mechanical gripping units of the
tensile tester at their ends, leaving a 4-cm gauge length for
mechanical loading. Load–deformation data were recorded
at a deforming speed of 0.05 cm/s, and the stress-strain
curve of the nanofibrous structure was constructed from the
load-deformation curve.
Scanning electron microscopy
Electrospun nanofibrous structures were sputter coated
with gold (Denton Desk-1 Sputter Coater), and their morphologies before and after cell seeding were observed by
scanning electron microscopy (SEM, Amray 3000) at an accelerating voltage of 10 or 20 kV.
Cell proliferation assay
Electrospun nanofibrous structures were sterilized by UV
light for 6 h per side. Utilizing a protocol approved by the
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Thomas Jefferson University Institutional Review Board,
human bone-marrow-derived mesenchymal stem cells
(hMSCs) were isolated from healthy donors with the patients’ consent.34 A 5–6-mL aliquot of bone marrow was
harvested from the iliac crest and spun at 600× g for 6 min
to sediment the red blood cells. The upper layer of the marrow suspension was added to 150-cm2 tissue culture flasks
(Corning Glass Works, Corning, NY) containing Dulbecco’s
modified Eagle’s medium (DMEM, BioWhittaker, Walkersville, MD), 10% fetal bovine serum (FBS, Premium Select,
Atlanta Biologicals, Inc., Atlanta, GA), and antibiotics (50
␮g/mL of streptomycin, 50 IU of penicillin/mL; Cellgro,
Herndon, VA).
The medium was replaced every three days and cultures
were maintained in a tissue culture incubator at 37°C with
5% CO2. After two passages, cells reached confluence and
were removed by trypsin treatment, counted, and seeded on
1-mm thick electrospun nanofibrous scaffolds (1 cm2) at a
density of 25,000 cells/cm2. The cellular constructs were
maintained in an incubator at 37°C, and the cell culture medium was changed every 3 days. After 1, 3, 5, 7, and 10 days,
the cellular constructs were harvested, washed with PBS to
remove non-adherent cells, then exposed to esterified dye,
2⬘, 7⬘-bis-(2-carboxyethyl) -5-(and-6)-carboxyfluorescein,
acetoxymethyl ester (BCECF-AM; Molecular Probes, Eugene, OR) for 30 min. The intensity of fluorescent dye
yielded during the cell-BCECF-AM incorporation was proportional to cell number.
Cell numbers at different time points were indicated indirectly by the relative fluorescence units (RFU) obtained
from readings from a SPECTRAFLUOR Plus fluorometer
(Tecan, Research Triangle Park, NC). The RFU readings of
three known cell numbers, 50,000, 100,000, and 200,000, were
used to create a standard curve to convert the RFU readings
to cell numbers.
Cell-matrix interaction
BALB/c C7 mouse fibroblast cells were obtained from
American Type Culture Collection (ATCC; Arlington, VA).
The cells were plated in monolayer in 75-cm2 tissue culture
flasks (Corning Glass Works, Corning, NY) and cultured to
confluence in cell culture medium consisting of Dulbecco’s
modified Eagle’s medium (DMEM; BioWhittaker, Walkersville, MD), 10% fetal bovine serum (FBS, Premium Select,
Atlanta Biologicals, Inc., Atlanta, GA), and antibiotics (50
␮g/mL of streptomycin, 50 IU of penicillin/mL; Cellgro,
Herndon, VA).
The medium was replaced every 3 days and cultures were
maintained in a tissue culture incubator at 37°C with 5%
CO2. A population of 50,000 BALB/c C7 fibroblast cells was
seeded on a scaffold and grown in DMEM medium with
10% FBS for 7 days. Cellular constructs were harvested at
days 1, 3, and 7, fixed with 4% glutaraldehyde for 1 h at
room temperature, dehydrated through a series of graded
alcohol solutions, and then air-dried overnight. Dry cellular
constructs were sputter coated and observed by SEM at an
accelerating voltage of 20 kV.
616
LI ET AL.
Statistical analysis
Values were expressed as means ± standard deviations.
Statistical differences were determined by Student’s twotailed t test.
RESULTS
ated. A representative stress–strain curve was constructed from the load–deformation curve and is illustrated in Figure 4. The tensile modulus of the structure
was 323.15 MPa. The ultimate tensile stress of the
structure was 22.67 MPa, and the ultimate strain was
95.8%. Since the composed fibers were randomly oriented within the nanofibrous structures, mechanical
anisotropy is expected on the X-Y plane of the structure.
Morphology of electrospun nanofibrous structures
SEM micrographs of electrospun nanofibrous structures revealed that the structure was composed of randomly oriented fibers the diameters of which ranged
from 500 to 800 nm (Fig. 2). Three-dimensional pores
formed between fibers were interconnected and distributed throughout the structure.
Porosimetry of electrospun nanofibrous structures
Assessment of structural pore properties was determined with the use of a mercury porosimeter. The
porosity of electrospun nanofibrous structure was
91.63%, indicating it was a highly porous structure.
The total pore volume was 9.69 mL/g, and the total
pore area was 23.54 m2/g. A representative plot of
pore diameter distribution, illustrated in Figure 3, indicates that pore diameters ranged broadly from 2 to
465 ␮m.
Tensile properties of electrospun
nanofibrous structures
Mechanical properties, such as tensile modulus, ultimate tensile stress, and ultimate strain, were evalu-
Cell proliferation assay
A cell proliferation assay was used to indicate the
number of living cells in a scaffold. Cell number was
converted from relative fluorescence units (RFU) by
calibrating a set of control experiments with a known
cell number. Following the seeding of 25,000 cells/cm2
on the nanofibrous scaffold, the cell number increased
with time, as illustrated in Figure 5, and reached a
plateau after day 7, showing a fivefold increase in cell
population during the 10-day culture period.
Cell–matrix interaction
Cell morphology and the interaction between cells
and nanofibers were studied in vitro for 7 days.
BALB/c C7 mouse fibroblasts adhered and spread on
the surface of the PLGA fiber network (Fig. 6) and had
started to migrate through the pores and to grow under layers of the fiber network at day 3 [Fig. 7(A)].
These fibroblasts interacted and integrated well with
the surrounding fibers [Fig. 7(B)]. SEM micrographs
showed that the development of cell growth was
guided by the fiber architecture. Cells grew in the direction of fiber orientation, forming a threedimensional and multicellular network according to
the architecture of the nanofibrous structure [Fig.
7(C,D)].
DISCUSSION
Figure 2. SEM micrograph (original magnification ×1500)
of the electrospun PLGA nanofibrous structure composed of
randomly oriented ultra-fine fibers. Bar, 10 ␮m.
The electrospinning process provides a promising
means for creating a tissue-engineered scaffold. Since
the electrostatic spinning technique was first patented
in 1934,35 many of its applications have been studied
in different engineering areas. In this investigation, the
nanofiber-based scaffold produced by the electrospinning process is introduced for the application of tissue
engineering. To date, many polymers have been electrospun, including polyethylene oxide (PEO), 32
acrylic,23 nylon,36 polyethylene glycol (PEG), polyac-
ELECTROSPUN NANOFIBROUS SCAFFOLD
617
Figure 3. Representative plot of pore diameter distribution. Each log differential intrusion value indicates the relative
quantity of pores of a specific diameter.
rylonitrile (PAN), polyethylene terephthalate (PET),37
poly(p-phenylene terephthalamide) (PPTA).38
PLGA is the most commonly used biodegradable
polymer in tissue engineering. The success of PLGA
electrospinning expands the applications of electrospinning technology. Other biodegradable polymers,
such as polylactic acid (PLA), polyglycolic acid (PGA),
and polyphosphazene, currently are under investigation in our laboratory.
A significant advantage of the electrospinning process is the ability to produce a nonwoven, nanometerscale fibrous structure. This feature is exploited to produce the PLGA nanofiber-based scaffold, which has
two unique features that make it well suited for tissue
engineering. First, most natural ECMs are composed
of randomly oriented collagens of nanometer-scale diameters. The morphology and architecture of the electrospun structure is similar to those of some natural
ECM.39 Second, the building components of the scaffold are made of biodegradable PLGA polymer. The
architecture of the scaffold is dynamically changed as
the polymer fibers are hydrolyzed and degraded over
time, which allows the residing cells to build up their
own ECM.
Pores in a tissue-engineered scaffold make up the
space in which cells reside. Pore properties, such as
porosity, dimension, and volume, are parameters directly related to the success of a scaffold. A PLGA
electrospun nanofibrous scaffold has more than 90%
porosity, which indicates a highly porous structure.
High porosity provides more structural space for cell
accommodation and makes the exchange of nutrient
and metabolic waste between a scaffold and environment more efficient. These characteristics are fundamental criteria for a successful tissue-engineered scaffold.40
This structure also features a wide variety of pore
diameters, varying from a few micrometers to several
hundred micrometers. When nanofibers are created in
the electrostatic field, they are deposited randomly on
the target layer by layer, and various diameters of
pores (distances between fibers) are formed. However,
the majority of pore diameters is limited to the range
of 25 to 100 ␮m. This scale of pore diameter is adequate for cell migration of most cell types, such as
fibroblasts, whose diameters are about 10–15 ␮m.
Previous investigations have mentioned the relationship between pore diameter and tissue ingrowth.
For vascular grafts, the effective pore diameters for
cell ingrowth are between 20 and 60 ␮m while for
bone ingrowth 75- to 150-␮m pore-diameters are required.20 Moreover, a recent study suggests that bone
ingrowth does not require pore diameters larger than
100 ␮m.41
Pores in an electrospun structure are formed by differently oriented fibers lying loosely upon each other.
When cells perform amoeboid movement to migrate
through the pores, they can push the surrounding fibers aside to expand the hole as the small fibers offer
little resistance to cell movement. This dynamic archi-
Figure 4. Stress–strain curve of three electrospun nanofibrous structures.
618
Figure 5. Proliferation of human bone-marrow-derived
mesenchymal stem cells (hMSCs) seeded on electrospun
nanofibrous structures. Quadruplicate samples at each time
point were measured, and the cell number at day 10 had a
fivefold increase compared with that at day 1 (p < 0.005).
tecture provides cells with an opportunity to optimally adjust the pore diameter and grow into the scaffold even though some pores are relatively small.
Therefore, some pores with smaller diameters in this
structure may not hinder cell migration. Such a hypothesis of dynamic cell–scaffold interaction needs to
be further investigated.
The mechanical property of a scaffold is another
important aspect of its design. The purpose of a scaffold is not only for providing a surface for cell residence but also for maintaining mechanical stability at
the defect site of the host.42 A well-designed tissueengineered scaffold has to meet two mechanical requirements to be effective. The scaffold must retain
structural integrity and stability when a physician
handles and implants it into the defect site of the host.
And after surgery, the structure at the implant site
must provide sufficient biomechanical support during
the process of tissue regeneration and structure degradation.
Some biopolymers, such as agarose and alginate,
have been shown to promote cell growth and differentiation, but their weak mechanical properties limit
LI ET AL.
clinical use. On the other hand, the synthetic polymerbased scaffolds, such as those composed of PLGA
polymer, have more effective mechanical properties.
The mechanical properties of an engineered scaffold
are dominated by intrinsic factors, such as the chemistry of the material, and by extrinsic factors, such as
the dimension or architectural arrangement of the
building blocks.
An electrospun nanofibrous structure is composed
of randomly oriented nanometer-scale PLGA fibers.
Its mechanical properties are determined by PLGA
molecular chains and the structural arrangement of
the nanofibers. Previous study has shown that the
electrospun nanofibrous structure has better mechanical properties than do structures composed of larger
diameter fibers.27 In addition, Amornsakchai et al.
found that the strength of fiber decreases as the fiber
dimensions increase.43 This novel electrospun nanofibrous structure will be stronger than the existing fiberbased structures composed of approximately 20 ␮m in
diameter fibers.
Compared with the results of experiments conducted by Kempson and Edwards,44,45 the mechanical
properties of the structure are comparable to those of
human cartilage and skin. The comparison of tensile
modulus, ultimate tensile stress, and ultimate strain
among electrospun structures, cartilage, and skin is
shown in Table I. Based on the morphology and mechanics of the scaffold, it is suggested that the electrospun nanofibrous structure would be suitable as a skin
or cartilage substitute.
In the cell proliferation assessment, the electrospun
nanofibrous structure was found to promote cell
growth. This novel structure provides a high level of
surface area for cells to attach to due to its threedimensional feature and its high surface area-tovolume ratio. It is important that a tissue-engineered
scaffold be loaded with the proper cell density without compromising the size of the scaffold. After 7 days
Figure 6. SEM micrographs (original magnification ×500) of BALB/c C7 mouse fibroblasts on an electrospun nanofibrous
structure at (A) 3 days of culture and (B) 7 days of culture. Bar, 10 ␮m.
ELECTROSPUN NANOFIBROUS SCAFFOLD
619
Figure 7. SEM micrographs of the interaction between cells and an electrospun nanofibrous structure after 3 days of culture:
(A) original magnification ×1500; (B) original magnification ×2500; (C) original magnification ×2500; and (D) original magnification ×3000. Bar, 10 ␮m.
of culture, cell number reaches a plateau, possibly because cells may have occupied all available spaces in
the prepared scaffold. Another possibility is that cell
number increase is interfered with by the effects of
PLGA fiber degradation. As degradation of PLGA occurs, the local environment inside the scaffold becomes acidic, which retards cell growth.
The electrospun nanofibrous scaffold provides a
three-dimensional structure for cell attachment,
growth, and migration. SEM micrographs reveal the
micro observation on the cellular response to this
structure. In this study, cells seeded on the nanofibrous scaffold are found to have appropriate interactions with their environment based on the following
observations. First, the cells maintain a normal phenotypic shape, suggesting that cells function biologically
within this structure. Second, the cells favor this structure, so they adhere onto the fibers and proliferate on
the nanofibrous network, packing the structure surface with cells after 7 days of culture. Third, these cells
crosslink the nanofibers and integrate with the surrounding fibers to form a three-dimensional cellular
network. Overall, this evidence indicates that nanofi-
brous structures positively promote cell–matrix and
cell–cell interactions, thus regulating signals from matrix and neighbor cells to induce the seeded cells in
this structure to express the phenotypic shape.
CONCLUSIONS
This study describes a homogenous, threedimensional structure consisting of ultrafine fibers
produced by the electrospinning process. The archiTABLE I
Comparison of Tensile Property
Electrospun Structure
Cartilage
Tensile modulus
(MPa)
Ultimate tensile stress
(MPa)
Ultimate tensile strain
(%)
Skin
323
130
15–150
23
19
5–30
96
20–120
35–115
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LI ET AL.
tecture of the structure is similar to that of natural
extracellular matrix, indicating that the electrospun
nanofibrous scaffold is suitable as a tissue substitute.
These nanofiber-based scaffolds are characterized
by a wide range of pore size distribution, high porosity, and high surface area-to-volume ratio, which are
favorable parameters for cell attachment, growth, and
proliferation. The structures provide effective mechanical properties suitable for soft tissue, such as skin
or cartilage. This study provides a basis for future
optimization of electrospun nanofibrous scaffold for
tissue-engineering applications. Applying this structure as a substitute for a specific tissue, such as cartilage, will be a focus of our future investigations.
The authors thank John Mckelvie for assistance in preparation of the electrospinning apparatus and Dr. Keith G.
Danielson for reviewing the manuscript.
References
1.
2.
3.
4.
5.
6.
7.
8.
9.
10.
11.
12.
13.
Langer R, Vacanti JP. Tissue engineering. Science 1993;260:920–
926.
Thomson RC, Yaszemski MJ, Mikos AG. Polymer scaffold processing. In: Lanza RP, Langer R, Chick WL, editors. Principles
of tissue engineering. Austin, TX: R.G. Landes; 1997. p 263–271.
Reddi AH. Morphogenesis and tissue engineering of bone and
cartilage: Inductive signals, stem cells, and biomimetic biomaterials. Tissue Eng 2000;6:351–359.
Alberts B, Bray D, Lewis J, Raff M, Roberts K, Watson JD.
Molecular biology of the cell. New York: Garland; 1994. p 971–
995.
Wintermantel E, Mayer J, Blum J, Eckert KL, Lüscher P,
Mathey M. Tissue engineering scaffolds using superstructures.
Biomaterials 1996;17:83–91.
Boyan BD, Hummert TW, Dean DD, Schwartz Z. Role of material surfaces in regulating bone and cartilage cell response.
Biomaterials 1996;17:137–146.
Dahlberg L, Kreicbergs A. Demineralized allogenic bone matrix for cartilage repair. J Orthop Res 1991;9:11–19.
Watt FM, Dudhia J. Prolonged expression of differentiated
phenotype by chondrocytes cultured at low density on a composite substrate of collagen and agarose that restricts cell
spreading. Differentiation 1988;38:140–147.
Ponticiello MS, Schinagl RM, Kadiyala S, Barry FP. Gelatinbased resorbable sponge as a carrier matrix for human mesenchymal stem cells in cartilage regeneration therapy. J Biomed
Mater Res 2000;52:246–255.
Allemann F, Mizuno S, Eid K, Yates KE, Zaleske D, Glowacki
J. Effects of hyaluronan on engineered articular cartilage extracellular matrix gene expression in 3-dimensional collagen scaffolds. J Biomed Mater Res 2001;55:13–19.
Bradham DM, Horton WE. In vivo cartilage formation from
growth factor modulated articular chondrocytes. Clin Orthop
1998;352:239–249.
Bonaventure J, Kadhom N, Cohen-Solal L, Ng KH, Bourguignon J, Lasselin C. Reexpression of cartilage-specific genes
by dedifferentiated human articular chondrocytes cultured in
alginate beads. Exp Cell Res 1994;212:97–104.
Agrawal CM, Ray RB. Biodegradable polymeric scaffolds for
musculoskeletal tissue engineering. J Biomed Mater Res 2001;
55:141–150.
14. Behravesh E, Yasko AW, Engel PS, Mikos AG. Synthetic biodegradable polymers for orthopaedic applications. Clin Orthop 1999;367S:S118–S125.
15. Athanasiou KA, Niederauer GG, Agrawal CC. Sterilization,
toxicity, biocompatibility, and clinical applications of polylactic acid/polyglycolic acid copolymer. Biomaterials 1996;17:93–
102.
16. Sanders JE, Stiles CE, Hayes CL. Tissue response to singlepolymer fibers of varying diameters: Evaluation of fibrous encapsulation and macrophage density. J Biomed Mater Res
2000;52:231–237.
17. Wan H, Williams RL, Doherty PJ, Williams DF. A study of cell
behaviour on the surfaces of multifilament materials. J Mater
Sci: Mater Med 1997;8:45–51.
18. Flemming RG, Murphy CJ, Abrams GA, Goodman SL, Nealey
PF. Effects of synthetic micro- and nano-structured surfaces on
cell behavior. Biomaterials 1999;20:573–588.
19. Green AM, Jansen JA, van der Waerden JPCM, von Recum AF.
Fibroblast response to microtextured silicone surfaces: Texture
orientation into or out of the surface. J Biomed Mater Res 1994;
28:647–653.
20. von Recum AF, Shannon CE, Cannon CE, Long KJ, van Kooten
TG. Surface roughness, porosity, and texture as modifiers of
cellular adhesion. Tissue Eng 1996;2:241–253.
21. Tan W, Krishnaraj R, Desai TA. Evaluation of nanostructured
composite collagen–chitosan matrices for tissue engineering.
Tissue Eng 2001;7:203–210.
22. Kadler KE, Holmes DF, Trotter JA, Chapman JA. Collagen fibril formation. Biochem J 1996;316:1–11.
23. Baumgarten PK. Electrostatic spinning of acrylic microfibers. J
Colloid Interface Sci 1971;36:71–79.
24. Doshi J, Reneker DH. Electrospinning process and application
of electrospun fibers. J Electrostatics 1995;35:151–160.
25. Reneker DH, Chun I. Nanometre diameter fibres of polymer,
produced by electrospinning. Nanotechnology 1996;7:216–223.
26. Fong H, Chun I, Reneker DH. Beaded nanofibers formed during electrospinning. Polymer 1999;40:4585–4592.
27. Kim J, Reneker DH. Mechanical properties of composites using
ultrafine electrospun fibers. Polym Composites 1999;20:124–
131.
28. Chun I, Reneker DH, Fong H, Fang X. Carbon nanofibers from
polyacrylonitrile and mesophase pitch. J Adv Mater1999;31:36–
41.
29. Ko FK, Laurencin CT, Borden MD, Reneker DH. The dynamics
of cell–fiber architecture interaction. Proc Ann Mtg Biomater
Res Soc 1998;Apr: 11.
30. Ko FK, Borden MD, Laurencin CT. The role of fiber architecture in biocomposites: The tissue engineering approach. Proc
13th Intl Conf Composite Mater 2001;Jul: Abstract No. 1428.
31. Fertala A, Han WB, Ko FK. Mapping critical sites in collagen II
for rational design of gene-engineered proteins for cellsupporting materials. J Biomed Mater Res 2001;57:48–58.
32. Jaeger R, Bergshoef MM, Batlle CM, Schönherr H, Vancso GJ.
Electrospinning of ultra-thin polymer fibers. Macromol Symp
1998;127:141–150.
33. Caplan AI. Mesenchymal stem cells. J Orthop Res 1991;9:641–
650.
34. Caterson EJ, Nesti LJ, Li WJ, Danielson KG, Albert TJ, Vaccaro
AR, Tuan RS. Three-dimensional cartilage formation by bone
marrow-derived cells seeded in polylactide/alginate amalgam.
J Biomed Mater Res. Forthcoming.
35. Formhals A. Process and apparatus for preparing artificial
threads. US patent no. 1,975,504; 1934.
36. Gibson PW, Schreuder-Gibson HL, Rivin D. Electrospun fiber
mats: Transport properties. AIChE J 1999;45:190–195.
37. Warner SB, Buer A, Ugbolue SC, Rutledge GC, Shin MY. A
ELECTROSPUN NANOFIBROUS SCAFFOLD
fundamental investigation of the formation and properties of
electrospun fibers. Natl Textile Center Ann Rep 1998;Nov:83–
90.
38. Srinivasan G, Reneker DH. Structure and morphology of small
diameter electrospun aramid fibers. Polym Int 1995;36:195–201.
39. Nishido T, Yasumoto K, Otori T, Desaki J. The network structure of corneal fibroblasts in the rat as revealed by scanning
electron microscopy. Invest Ophthalmol Vis Sci 1988;29:1887–
1990.
40. Ma PX, Choi JW. Biodegradable polymer scaffolds with welldefined interconnected spherical pore network. Tissue Eng
2001;7:23–33.
41. Itälä AI, Ylänen HO, Ekholm C, Karlsson KH, Aro HT. Pore
diameter of more than 100 ␮m is not requisite for bone in-
621
42.
43.
44.
45.
growth in rabbits. J Biomed Mater Res, Appl Biomater 2001;
58:679–683.
Hutmacher DW. Scaffold in tissue engineering bone and cartilage. Biomaterials 2000;20:2529–2543.
Amornsakchai T, Cansfield DLM, Jawad SA, Pollard G, Ward
IM. The relation between filament diameter and fracture
strength for ultra-high-modulus polyethylene fibres. J Mater
Sci 1993;28:1689–1698.
Kempson GE, Muir H, Pollard C, Tuke M. The tensile properties of the cartilage of human femoral condyles related to the
content of collagen and glycosaminoglycans. Biochim Biophys
Acta 1973;297:456–472.
Edwards C, Marks R. Evaluation of biomechanical properties
of human skin. Clin Dermatol 1995;13:375–380.