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Materials Science & Engineering C
journal homepage: www.elsevier.com/locate/msec
Review
Ultrathin polymer fibers hybridized with bioactive ceramics: A review on
fundamental pathways of electrospinning towards bone regeneration
Filipe V. Ferreira a, Caio G. Otoni b, João H. Lopes c, Lucas P. de Souza d, Lucia H.I. Mei a,
Liliane M.F. Lona a, Karen Lozano e, Anderson O. Lobo f, *, Luiz H.C. Mattoso g, **
a
School of Chemical Engineering, University of Campinas (UNICAMP), Campinas, SP, Brazil
epartment of Materials Engineering ( EMa), e eral University of São Carlos (U SCar), São Carlos, SP, Brazil
epartment of Chemistry, ivision of n amental Sciences (IE ), echnological Instit te of Aerona tics (I A), São ose os Campos, SP, Brazil
d
College of Engineering an Physical Sciences, Aston Instit te of Materials esearch, Aston University, Birmingham, U
e
epartment of Mechanical Engineering, he University of e as io ran e alley, E in rg,
, USA
f
Inter isciplinary a oratory for A vance Materials, BioMat a , Materials Science an Engineering ra ate Program, e eral University of Pia i, eresina, PI, Brazil
g
Nanotechnology National a oratory for Agric lt re ( NNA), Em rapa Instr mentation, São Carlos, SP, Brazil
b
c
A R T I C L E I N F O
A B S T R A C T
ey or s
Bioactive glass
Hydroxyapatite
Polymer nanofibers
Biomedical application
Tissue engineering
Electrospinning
Electrospun ultrathin polymer fibers hybridized with bioactive ceramics find use in many biomedical applications due to their unique and versatile abilities to modulate structure–performance relationships at the nano–bio
interface. These organic–inorganic hybrid fibers present synergies that are otherwise rare, even when the precursors are used individually, such as bioactivity in polymers and stiffness–toughness balance in bioactive ceramics. Despite these unique advantages, a comprehensive and timely review on this important topic is still
missing. Herein we describe the most recent and relevant developments on electrospun ultrathin polymer fibers
hybridized with bioactive ceramics, with emphasis on bone tissue regeneration. This review addresses the
preparation of bioactive ceramics, particularly (nano) hydroxyapatite (HA; nHA) and bioactive glass (BG), which
stand out as the ceramics of interest for bone regeneration. The anatomy and mechanical properties of bone as
well as fundamental tissue–scaffold interaction mechanisms are covered. The process–structure–property relationships of electrospun ultrathin fibers are discussed in detail from a technical standpoint, as well as fabrication strategies, process variables, characterization methods, and biological requirements (in vitro and in vivo
performances). Finally, we highlight the major challenges and outline perspectives to pave the route for the nextgeneration hybrid materials for bone tissue engineering.
1. Introduction
The preparation of materials for tissue engineering has gained notoriety over the last years [1–4]. Specifically, in bone tissue engineering,
multidisciplinary research groups have focused their efforts on the
fabrication of scaffolds that allow efficient regeneration of lost or nonviable tissue [5–8]. In this scenario, over the last three decades, scientists have been designing ‘bio-inert’ materials for bone tissue engineering [9,1 ]; more recently, efforts have shifted to the preparation of
‘bioactive’ materials [11,1 ]. In contrast to ‘bio-inert’ materials that are
mainly used as substitutes (e g platinum and its alloys), bioactive
materials have garnered interest due to their ability to interact with
body cells, facilitating therefore the regeneration of tissues instead of
replacement [1 ]. Several bioactive materials have been investigated,
particularly bioactive ceramics such as hydroxyapatite (HA) and
bioactive glass (BG). Both HA and BG can be highlighted as bioactive
materials for bone tissue engineering as they have demonstrated good
capacity to enhance bone regeneration [14,15], acting mainly in the
expression of osteogenic genes [1 ] and stimulation of angiogenesis
[1 –19]. However, these compounds, in bul , do not match the features
of bone tissue, namely porous morphology, three-dimensional shape
stability, and hard but not brittle mechanical behavior, altogether
* Correspondence to: A O. Lobo, LIMAV - Interdisciplinary Laboratory for Advanced Materials, BiomatLab, Materials Science and Engineering Graduate Program,
UFPI - Federal University of Piau , 4 49-55 Teresina, PI, Brazil.
** Correspondence to: L.H.C. Mattoso, Nanotechnology National Laboratory for Agriculture (LNNA), Embrapa Instrumenta ão, 1 5 -9 Sao Carlos, SP, Brazil.
E mail a resses lobo ufpi.edu.br (A.O. Lobo), luiz.mattoso embrapa.br (L.H.C. Mattoso).
https: doi.org 1 .1 1
Received
October
j.msec.
.11185
; Received in revised form 1 December
; Accepted
December
erreira et al
Table 1
Typical properties of the selected polymers used to prepare ultrathin fibers
containing bioactive ceramics for bone tissue engineering.
Polymer
Synthesis*
Properties**
Ref.
PBAT
Polycondensation
[
, 4,
]
PLA
Ring-opening polymerization
and polycondensation
[
, 5,
, 8]
PCL
Ring-opening polymerization
PHB
Biotechnological routes
PHBV
Biotechnological routes
Tm = 11 –1 ◦ C
Tg = − ◦ C
E = – MPa
σ = – MPa
ε≥
Non-biodegradable
in the body
Tm = 1 –1 8 ◦ C
Tg = 4 ◦ C
E = . 5– .5 GPa
σ = 5 MPa
ε=5
Biodegradable in the
body
Tm = 58– ◦ C
Tg = − ◦ C
E = . 1– .44 GPa
σ = 4– 85 MPa
ε = –1
Biodegradable in the
body
Tm = 1 ◦ C
Tg = ◦ C
σ = 4 MPa
ε=5
Biodegradable in the
body
Tm = 145 ◦ C
Tg = −1 ◦ C
σ=
MPa
ε=5
Biodegradable in the
body
[11, 9]
[4 ]
[4 ]
*
e report the most used methods.
Melting temperature (Tm), glass transition temperature (Tg),
modulus (E), tensile strength (σ), and elongation at brea (ε).
**
oung’s
limiting their use in bone regeneration.
Several authors have developed electrospun bioactive ultrathin
nanostructured fibers by combining bioactive ceramics with polymer
systems [ , 1]. These fiber membranes mimic the ideal basic premises
for bone tissue engineering, namely: biocompatibility, porous structure,
surface roughness (suitable for cell attachment, migration, and proliferation), mechanical strength, and osteoinductive activity [ , ].
Despite the aforementioned promising features, many challenges remain
in this field, mostly related to the clinical use of these materials and their
production at an industrial scale [ 4, 5]. Strategies to overcome such
hurdles lie at every stage throughout the preparation route of these
emerging biomaterials, including synthesis and properties of the
bioactive ceramics, as well as fabrication and in vitro and in vivo performance s of polymer bioactive ceramic composite or hybrid fibers
[ ].
Herein, we provide an overview of the current state-of-the-art on
electrospun polymer-based ultrathin fibers filled with bioactive ceramics for bone tissue engineering. The synthesis protocols, ey properties, and potential applications of the most common bioactive
ceramics, for bone tissue regeneration purposes, are highlighted and
discussed in detail. e also address the bone tissue itself focusing on the
biological requirements of its scaffolds. The most important aspects of
ultrathin polymer fibers filled with bioactive ceramics and their applications in bone tissue engineering are covered, including production,
characterization, and in vitro and in vivo performances. Finally, we
identify and suggest means of overcoming the main challenges on the
use of these biomaterials, as well as outline perspectives for their clinical
usage.
2. Polymer matrix
A number of natural and synthetic polymers has been used as
matrices or binding agents to the bioactive ceramics, among which the
most studied ones are poly(butylene adipate-co-terephthalate) (PBAT),
poly(lactic acid) (PLA), poly(ε-caprolactone) (PCL), poly(DL-lactide)
(PDLLA), poly(lactic-co-glycolic acid) (PLGA), poly( -hydroxybutyrate)
(PHB), and poly( -hydroxybutyrate-co-hydroxyvalerate) (PHBV). The
synthesis and properties of these polymers can be found elsewhere
[ – 5]. The overall properties of these polymers are summarized in
Table 1.
PBAT is considered as one of the most attractive polymers to prepare
ultrathin fibers due to its easy processing and favorable mechanical
properties. In particular, highly exible PBAT-based materials can be
easily prepared using chloroform or toluene as solvents [41]. However,
the non-biodegradability of PBAT in the body represents a problem if the
implanted system needs to be degraded. PLA and PCL can be degraded
under physiological conditions (such as in the human body). In this
context, when the biodegradabilities of PLA and PCL are compared, the
degradation inetics of PCL is even slower, and therefore this polymer
has received greater attention for the development of long-term
implantable biomaterials. However, the widespread use of PLA and
PCL has been hindered by numerous mechanical shortcomings. For
example, PLA presents high strength but low toughness, whereas PCL
presents low oung’s modulus combined with low cell adhesion.
It is well nown that the two attributes “tough” and “strong” can be
mutually exclusive, but a good strength-toughness balance can be achieved along with acceptable cell adhesion [4 ]. Therefore, PBAT-, PLA-,
and PCL-based blends have been formulated to enhance toughness and
bioactive properties. These polymers have shown optimum miscibility
among themselves and in other polymers such as poly(vinyl chloride),
polycarbonates, nitrocellulose, and cellulose butyrate [4 ].
PHB and PHBV are some polyhydroxyal anoates that are also used to
prepare ultrathin fibers for bone tissue engineering. In general, the stiff,
brittle nature of these (co)polymers leads to ultrathin fibers that behave
as a hard-elastic material [44]. This behavior is more evident in fibers
prepared with PHB than in those with PHBV. Despite these mechanical
limitations, the biodegradability of both PHB and PHBV drives the usage
of these materials [4 ].
3. Bioactive ceramics
In the field of biomaterials, an important milestone occurred in the
late 19 s, which represented a paradigm shift with the genesis of a new
class of biomaterials, i e , bioactive materials [45]. The bioactivity is
related to the ability of a biomaterial to stimulate a beneficial response
through interactions with the host tissue [4 –5 ].
Recently, bioactive ceramics have attracted significant attention
owing to their excellent bioactivity, osteostimulatory, and biodegradable properties [51]. Here, we brie y report on the synthesis, properties,
and potential applications of the two main bioactive ceramics used as
fillers in scaffolds for bone tissue regeneration: BG and HA.
Bioactive glasses
BG, developed in the late 19 s by Professor Dr. Larry Hench of the
University of Florida in Gainesville, were the first materials with an
ability to form bonds directly with both hard and soft tissues and , as
such, stimulate bone growth (formation of a new tissue) [4 ,5 ]. The
melt-derived 45S5 Bioglass® (45SiO - 4.5Na O- 4.5CaO- P O5, in wt
) was the first composition of bioactive glass developed. The high
amounts of Na O and CaO ma e the surface of the material very reactive
in the physiological environment.
In the following decades, several compositions were extensively
studied and their bioactivity evaluated for biomedical applications,
many of which became commercial products, such as those in the
erreira et al
Fig. 1. Schematics of the chemical and biological transformations that occur on the surface of the bioactive glass when placed in a physiological environment.
Bioglass® family [4 ]. The ability of BG to bind to the bone tissue has
been attributed to the formation of a layer of hydroxycarbonate-apatite
(HCA) on the glass surface in contact with the physiological environment [5 ,54]. The chemical composition and structure of the HCA layer
is similar to that of the bone, which provides a strong bond at the
interface [48,5 ]. The mechanism that explains the bioactivity of the
BG, proposed by Hench et al., was based on studies of soda-lime glass
corrosion and confirmed by infrared spectroscopy [55]. This comprises a
partial dissolution of the vitreous networ in the presence of body uid,
which is governed by sequential chemical reactions and biological
events, as illustrated in Fig. 1.
Brie y, the BG initially loses sodium and calcium ions by exchange
with H O+ ions of the physiological environment. This leaching promotes a local pH increase, leading to the brea down of the chemical
bonds present in the Si–O–Si groups , consequently releasing the soluble silica, as shown in Eq. (1) [5 ]:
Si–O–Na+ + H+ + OH− → Si⎯OH+ + Na+ (aq) + OH−
(1)
This dissolution leads to the formation of amorphous SiO -rich gel
layer on the glass surface, which subsequently forms a biologically
active amorphous film of calcium phosphate (CaO⎯P O5) through the
incorporation of Ca + and PO4− species from body uids. Another
incorporation then occurs (in this case OH− and CO −), which evolves
towards the polycrystalline phase of HCA [48,55,5 ,58].
The biological events that culminate in the formation of the interfacial bond between the BG and the host tissue occur concurrently with
the chemical reactions that transform the glass surface into a cellfriendly environment [5 ,59].
BG can be used in the treatment of gastric ulcers [ ], s in wounds
and burns [ 1], bone regeneration [ ], soft tissue repair [ ], and as
drug carriers for cancer treatment [ 4]. The observed exceptional
behavior of BG in the field of biomaterials has been attributed mostly to
the dissolution products released from the glass surface (e g Ca, Si, and P
ions), which stimulate the genes of cells towards a path of tissue
regeneration. ynos et al. [ 5, ] showed that the controlled release of
these ions from the BG can control the cell cycle of osteoblasts. The
osteoprogenitor cells colonize the BG surface, receiving the correct
chemical stimuli from their environment to enter the active segments of
the cell cycle. This behavior leads to cell division (mitosis), subsequently
undergoing osteogenic differentiation, and eventually promoting bone
formation (osteogenesis). The chemical stimulus has been related to the
presence of the Si and Ca ions at the solution cell interface, which in
critical concentrations in uence the expression of a series of genes
responsible for osteogenesis. Gene families that are upregulated are
related to the relevant cell cycle segments, cell proliferation, and cell
differentiation [1 , 5– 8].
Other inorganic species including boron, copper, cobalt, silver, zinc,
and strontium may play an important role in ion therapy [ 9]. Souza and
colleagues [ ] reported a comprehensive study on the bioactive
properties of Nb-substituted silicate glass and showed that the presence
of niobium species may be associated with an increase of osteostimulation in the BG. The presence of these species in appropriate quantities
induces intracellular and extracellular responses at the genetic level
in uencing important biological processes such as bone healing
[5 , 8, 1, ]. The bioactivity depends on glass composition as well as
its structural characteristics (short- and long-range ordering) [ ].
The use of BG as delivery systems for ion therapy is not restricted to
applications in bone regeneration. Doping BG with gallium ions has
shown significant efficacy in aiding in the treatment of bone cancer, the
current treatment of which consisting of surgical harvesting of as much
cancer tissue as possible and subsequent chemotherapy, and in some
cases radiotherapy. An efficient biomaterial to aid in the treatment of
bone cancer should be able to (i) aid in the regeneration of bone deficiency created by the surgery; and (ii) ill any remaining cancer cell in
the affected site [ 4]. Ga + is nown to compete with Fe + for transferrin bonding sites and its subsequent intracellular pathways in biological systems [ 4]. As cancerous cells present a significantly higher
number of transferrin receptors on their surfaces, they can upta e much
more Ga +. High intracellular levels of Ga + were shown to be toxic to
the cells and the treatment with gallium tends to arrest cancer cells to
apoptosis. On the other hand, in the normal cells the intracellular levels
of Ga + do not reach the toxic threshold because there are not enough
transferrin receptors on their surfaces. Rana and colleagues [ 4]
exposed human osteosarcoma cells (SAOS- ) and Normal Human Osteoblasts (NHOsts) to the dissolution products of silica-based bioactive
glasses containing 1, , or wt of Ga O for
h. They observed that
the composition containing
Ga O selectively illed bone cancer
cells not affecting the healthy cells while maintaining the same bioactive
properties of the original bioglass 45S5 composition, essential features
for the regeneration of post-surgical bone defect [ 4].
Generally, BG, in particular silica-based glasses, comprise high
amount s of Na O and CaO as well as relatively high CaO P O5 ratios
[58]. The preparation method is a preponderant factor with respect to
the microstructure of the BG [ ], which are commonly prepared by
melt-quenched or sol-gel methods [ 5]. The sol-gel route has some advantages over the melt-quenched method and allows a low temperature
erreira et al
Fig. 2. Schematic diagram of cell types associated with the bone tissue.
synthesis for BG with high silica contents, resulting also in systems with
higher specific surface areas. However, in some cases, it can be challenging to obtain glasses with high structural homogeneity, particularly
for complex glass compositions. Several studies have focused on developing strategies to control the microstructure of glasses obtained via solgel. Lopes et al. [ ] reported an efficient strategy for the synthesis of
58S BG with high structural homogeneity using a citric acid-assisted solgel.
y ro yapatite
Hydroxyapatite (HA) is a biocompatible and biodegradable ceramic
represented by the general formula Ca5(PO4) OH and unit cell formula
Ca1 (PO4) (OH) . This bioceramic can be prepared in multiple morphologies (nanowires, nanorods, microspheres, microsheets, etc ),
different sizes (including nano-HA; nHA) [ , ], and by several
methods (such as precipitation [ 8], electrospraying [ 9], mechanochemical synthesis [8 ], multiple emulsion [81], and others [8 ,8 ]).
Several characteristics have been reported about the important role of
HA applications as a biomaterial, such as size and crystal morphology,
dispersion, and dissociation delay [84–8 ]. These recent advances have
been discussed extensively in other review articles [8 ,8 ] and therefore
not repeated in detail here.
Due to the close similarities with the inorganic phase of bone, HA
shows broad applications in bone tissue engineering (bone repair,
coating filler of implants, etc ) [88]. This behavior is related to the
integration ability of HA to create quic bonds with neighboring bones
due to its chemical, structural, and morphological properties. In the case
of nano-structured HA, this integration ability is improved mainly due to
unique surface properties of this nano-sized material (high specific
surface area and increased roughness), which lead to a greater reactivity
and therefore better cell adhesion and cell-matrix interactions [89]. The
enhanced bone integration offers superior biosorption with greater
advantage in clinical applications [9 ]. However, the low mechanical
strength of (nano-)HA restricts its use [8 ]. Several authors have suggested the addition of these ceramic materials into polymer composites
for the development of biomedical materials with improved mechanical
and bone integration performances. This alternative may represent a
solution to presently observed problems and therefore commercial opportunities [91,9 ].
4. Biomaterial for bone tissue engineering
The development of bone substitutes requires ample understanding
of several aspects that comprise the bone tissue such as type of cells,
physiology, mechanical properties, composition and arrangement of its
extracellular matrix, and micro macro architectures.
Anatomy an mechanical properties of ones
Bone is a specialized form of connective tissue composed of cells and
a mineralized extracellular matrix. It provides structural support and
locomotion for the body as well as protection for internal organs [9 ].
Five cell types are associated with bone: osteoprogenitor cells, osteoblasts, bone-lining cells, osteocytes, and osteoclasts (Fig. ) [94,95].
Osteoprogenitor cells are derived from mesenchymal stem cells and
differentiate into osteoblasts through the in uence of CBFA1 transcription factor (RUN ) [9 ]. Osteoblasts drive bone formation, i e ,
synthesize and regulate bone deposition and mineralization [5]. Osteocytes are mature osteoblasts that are enclosed within the lacunae in the
mineralized matrix and aid in its adequate maintenance. Bone-lining
cells derive from osteoblasts and remain on the bone surface when no
active growth is occurring. Osteoclasts, in turn, are phagocytic cells that
derive from hemopoietic progenitor cells in the bone marrow. They
differentiate and mature into osteoclasts under control of the receptor
activator of nuclear factor appa-В ligand (RANKL) signaling mechanism and act in bone resorption, participating in the processes of bone
remodeling and blood calcium homeostasis [9 ,9 ].
The most distinct feature of the bone tissue is the presence of the
mineralized extracellular matrix that is responsible for its hardness and
provides support and protection capabilities. The microscopic arrangement of the bone matrix is well understood and has been extensively
described elsewhere [98–1 1]. Brie y, it comprises three major phases:
organic (ca
proteins), inorganic (ca
calcium phosphate in the
form of hydroxyapatite crystals [Ca1 (PO4) (OH) ]), and water (ca
1 ) [1 ]. The organic phase comprises 9
of collagenous proteins
(mainly type-I collagen) and 1
of non-collagenous proteins such as
proteoglycans, glycoproteins, vitamin K-dependent proteins, growth
factors, and cyto ines. The organic phase imparts toughness to bone and
supports the crystalline (inorganic) phase, while the strength comes
from the latter [1 ,1 4].
All of the components described above constitute microscopic
mineralized spicules named bone trabeculae. The three dimensional
erreira et al
Fig. 3. Schematics of a generic electrospinning apparatus used for the production of polymer-based, bioactive ceramic-filled ultrathin fibers, which involves (a) the
preparation of a polymer ceramic fiber-forming formulation, (b) a syringe pump, (c) a syringe, (d) a high-voltage generator, and (e) a grounded collector.
distribution of such trabeculae within the bone determines the mechanical properties of the different types of bones [1 5,1 ]. here
these trabeculae are compacted together, the bone is called cortical or
compact bone and its microarchitecture presents porosity between 5
and
, compressive strengths of 1 –
MPa, exural strengths of
1 5–19 MPa, tensile strengths of 5 –151 MPa, and elastic moduli of
1 –18 GPa. This type of bone tissue is most commonly encountered in
the shaft (also called diaphysis) of long bones, such as those of limbs,
fingers, and toes. On the other hand, where these trabeculae are more
dispersed, the bone is named cancellous, trabecular or spongy bone and
presents significantly higher porosity (between
and 9 ),
compressive strengths of –1 MPa, tensile strengths of 1– MPa, and
elastic moduli of .1– .5 GPa [1 5]. Trabecular bone is mostly located
in the extremities of long bones (called epiphysis), and within short and
irregular bones such as those of feet, face, and vertebrae [1 5,1 ].
Bone iomaterial
Scaffolds, cells, and growth factors are the three basic elements of
bone tissue engineering [1 ]. The regeneration and self-repairing capabilities of bone are well nown [1 ]. However, in some cases
(depending on the type and location of the fracture site), a fracture fails
to heal, leading to a non-union that requires the use of bone substitutes
for successful healing. The incidence of non-unions has been shown to be
higher in the 5–44-year age group, bringing significant financial implications since the overall costs per patient are estimated to range between 1,18 and
, 5 [9 ]. The increasing demand for bone tissue
reconstruction has motivated the development of new scaffolds
[1 ,1 8,1 9]. Bone graft substitutes can be used to repair damaged or
fractured tissues [1 9] as long as they satisfy biological requirements
such as biocompatibility (non-toxicity) and biodegradability as well as
present specific structural properties for optimized interaction with the
bone tissue [11 ].
I eal scaffol microarchitect re for one tiss e engineering
For ideal integration and support of bone regeneration, scaffolds
should exhibit similar microstructure of that aforementioned for native
bones. The microarchitecture of the scaffold, which includes its porosity,
pore size, and interconnectivity between the pores, as well as its surface
roughness are ey for effective traffic of nutrients and waste products, as
well as vascularization, tissue infiltration, and cellular adhesion [111].
Sufficiently large and interconnected pores (between 1 and 5 μm)
permit bone ingrowth due to larger availability of space, preventing
peripheral cellular growth and fibrosis. Also, the larger the surface area
of the scaffold, the greater is its contact with body uids, allowing
greater protein adsorption, therefore enhancing cell adhesion. However,
it is important to point out that excessive porosity will certainly reduce
mechanical performance of the scaffold. Thus, the ideal scaffold should
exhibit an equilibrium between these variables. Ideally, porosity should
match that of the bone intended to be replaced, thus nowing the
characteristics and distribution of the different types of bones in the
s eleton is essential for the effective design of bone scaffolds [11 ].
Surface roughness facilitates cell attachment, differentiation, and
maturation [11 ]. The mechanical stability of scaffolds within the body
assists in their adhesion to the surrounding bone tissue and supports the
cells [114]. The characteristics of the surface favors adsorption of adhesive proteins (e g , fibrin) onto which osteogenic cells attach, proliferate, and differentiate into osteoblasts, further producing bone that
integrates with the scaffold [5,115]. Electrospun polymer ultrathin fibers containing bioactive ceramics have recently been considered in
order to prepare materials that simultaneously combine suitable
biocompatibility, porosity, roughness, mechanical performance (structural integrity and exibility), and osteoinductivity.
Electrospun polymer ultrathin fibers have shown promising
biocompatibility, ensuring cell viability, adhesion, and proliferation
[11 ], besides offering high specific surface area. Regarding biodegradability, scaffolds ought to ideally remain inert or resorb at the same
rate at which new tissues are formed [11 –119]. Surface properties such
as porosity and roughness are important because these also affect the
formation of new tissue [111]. Specifically, porosity favors the infiltration of cells and blood vessels [11 ], whereas roughness facilitates
cell attachment, differentiation, and maturation [11 ].
5. Engineering ybrid electros un ultrat in olymer
nanointerface
bers at
Several authors have used bioactive ceramics (i e , HA and BG) as
fillers for various polymer matrices, acting as binders, into ultrathin fibers to improve cell compatibility and bone-forming process. Furthermore, this approach has demonstrated to overcome the consensual
major issue jeopardizing the use of bioactive ceramics as biomaterials :
brittleness [1 ]. These ultrathin fiber membranes also effectively
mimic the structure and function of natural bone [9 ]. Electrospun HAfilled ultrathin fibers, for instance, have been seen to feature a crystalline structure similar to that of bones [1 1], the BG induces the
erreira et al
Fig. 4. Schematic illustration of the parameters nown to dictate the outcome of the electrospinning process.
formation of bone-li e HA layers when in contact with body
[1 –1 4].
uids
n amentals of electrospinning
Electrospinning has been shown as a versatile laboratory method for
the preparation of polymer-based fibers with nano-sized diameters
[1 5]. These fibers combine the physical and chemical properties of the
precursor materials with the possibility of fine-tuning roughness and
porosity towards favored cell functions [1 –1 9]. The fabrication
process of nanofibers by electrospinning involves straightforward
physical and chemical principles: for the sa e of simplicity, it mainly
relies on the balance between electrostatic and surface tension forces. In
fact, other forces such as drag, gravity, Coulombic repulsion, and
viscoelastic forces may also be at play depending on the system. The
electrsostatic forces intentionally overcome the cohesive ones to elongate a liquid fiber-forming formulation (FFF) into a jet that emerges
from an electrically charged needle or nozzle (hereafter referred to as
nozzle) and is deposited onto a conductive collector [1 ,1 1]. In the
case of bioactive ceramic-filled ultrathin fibers, there is an additional
pre-spinning step involving the preparation of a suitable polymer
ceramic FFF, as illustrated by the generic electrospinning apparatus
schemed in Fig. . Subsequently, such FFF (Fig. a) is forced by a syringe
pump (Fig. b) through a nozzle (Fig. c) that is connected to a highvoltage supply (Fig. d). Finally, the material is deposited onto an
oppositely charged (usually grounded) collector (Fig. e). Depending on
several parameters, which have been extensively reviewed elsewhere
[1 ,1 ] and are detailed below within the context of tissue engineering, the deposited material can be a nanofibrous mat featuring
suitable characteristics for bone regeneration.
ers
Strategies for electrospinning polymer ase , ceramic lle
ltrathin
As mentioned above, electrospinning stands for a process in which a
liquid FFF is extruded within an electrically charged, confined channel,
then elongated and driven by a differential voltage towards a grounded
or oppositely charged collector, where fibers are deposited as
asi-
dried filaments as solvent evaporates.
Several parameters are nown to dictate the outcome of the electrospinning process (Fig. 4), including those related to the polymer itself
(e g , electrical conductivity, chemical structure, and molecular weight),
to the solutions or dispersions — hereafter referred to as FFFs (e g ,
viscoelasticity, surface tension, and ionic strength), to the processing
conditions (e g , voltage, nozzle shape and dimensions, and FFF feeding
rate), and to the surrounding environment (e g , temperature, relative
humidity, and pressure) [1 4,1 5].
iscosity of the sol tion
First, the FFF must be able to ow from a reservoir (typically a syringe) through a nozzle, normally forced by a syringe pump. At this
point, the rheological properties of the FFF stand out as ey parameters,
which in turn are in uenced by solid content, temperature, and injection
rate, to mention a few [1 ,1 ]. In this sense, for a specific polymer
solvent system, adjusting polymer and filler concentrations in FFFs has
been extensively addressed in optimization studies, as their viscoelastic
behavior is highly affected by the solid content [1 ,1 8]. In recent
studies focusing on bioactive ceramic-containing polymer fibers for
biomedical applications, polymer solutions ranging in concentration
from to 14 have been used to develop electrospun ultrathin fibers.
Given the typical shear thinning behavior of polymer solutions as well as
the dependence of viscoelasticity on temperature, these factors have also
been fine-tuned to achieve desired fiber morphology and yield.
The solvent – or mixture of solvents – to form solutions or to obtain a
dispersant medium in the case of bioactive ceramic-containing FFFs
plays an important role in FFF rheology, surface tension, evaporation
rate, and drying time (depending on its vapor pressure). Thus, solvent
dispersant volatility dictates not only the feasibility of electrospinning
certain FFFs, but also affects the yield and morphology of the resulting
fibers [1 ,1 ]. Highly volatile organic solvents have been mostly used
for the preparation of polymer-based bioactive ceramic-filled ultrathin
fibers, although efforts have been made towards the use of aqueous
dispersant media [1 9]. FFF feeding rate and nozzle-to-collector distance denote processing variables that are commonly adjusted depending mainly on solvent volatility.
erreira et al
Fig. 5. Schematic illustration of the Taylor cone formation: (i) formation of asi-spherical droplet, (ii) charges are induced in the polymer solution, (iii) formation of
Taylor cone and (iv) the “whipping instability” regime to form polymer fibers while the solvent evaporates.
Figure based on the wor of Ref. [14 ].
Table 2
Production of electrospun polymer-based bioactive ceramic-filled ultrathin fibers over the last five years: composition of fiber-forming formulations and the most
important processing parameters.
Organic matrix
Chitosan PEO
PBAT
PBAT
PBAT
PBAT PPy
PCL
PCL chitosan
PES
PHB
PHB PCL
PLGA
PLLA
PLLA
PU
PVOH
Inorganic filler
Content ( )
Solvent(s)
Voltage ( V)
Chloroform DMF
Chloroform DMF
DCM DMF
Chloroform DMF
TFE
Formic acetic acids
DMF
Chloroform
Chloroform DMF
HFIP
DCM DMF
Chloroform acetone
DMF THF
ater
1
14–18
a
BG
nHA
nHA
HA
nHA
45S5 BG
μBG and nBG
BG
nHA
58S BG
Mesoporous BG
nHA
BG
μHA and nHA
1
1
1
1
a
b
a
b
a
4a
b
b
5
1 a
14. a
.1b
15–
8a
a
1
18
15
8
4.5
14
15
1 .5
15
Feeding rate (mL h−1)
Distance (cm)
Ref.
.4
.8
.
1.5
1
5
.
.
1
1
1
1
15
1
18
1 .5
NR
[14 ]
[148]
[ ]
[149]
[15 ]
[1 ]
[151]
[15 ]
[15 ]
[154]
[155]
[1 ]
[14 ]
[145]
[1 9]
1
.4
.4
15
15
15
15
1
BG, bioactive glass; DCM, dichloromethane; DMF, dimethylformamide; HA, hydroxyapatite; HFIP, hexa uoroisopropanol; μBG, BG microparticles; μHA, HA microparticles; nBG, BG nanoparticles; nHA, HA nanoparticles; NR, non-reported; PBAT, poly(butylene adipate-co-terephthalate); PCL, poly(ε-caprolactone); PEO, poly
(ethylene oxide); PES, polyethersulphone; PHB, poly( -hydroxybutyrate); PLGA, poly(lactic-co-glycolic acid); PLLA, poly(L-lactic acid); PPy, polypyrrole; PU, polyurethane; PVOH, poly(vinyl alcohol); TFE, tri uoroethanol; THF, tetrahydrofuran. Distance refers to nozzle-to-collector.
a
eight percentage.
b
eight volume percentage.
oltage, aylor’s cone factor, an o rate
Another ey aspect is the applied voltage because it is responsible for
overcoming the cohesive forces and deforming the FFF into the so called
Taylor cone at the tip of the nozzle (Fig. 5). Specifically, asi-spherical
droplet are formed and electric charges are induced in the polymer solution (Fig. 5i–ii) due to the strong electrostatic field from the nozzle
(electrode) connected to the conductive collector screen (counter-electrode) [14 ]. hen the shear stresses due to repulsion induced from the
charges on the liquid surface is stronger than the surface tension of the
solution, a pendant drop with conical shape (Taylor cone — Fig. 5iii) is
formed [141]. The initial jet then drastically decreases in diameter until
it starts to bend, entering the “whipping instability” regime (Fig. 5iv)
[14 ]. Ultrathin polymer fibers are deposited onto the grounded collector while the solvent evaporates.
It has been demonstrated that the voltage not only serves as the
driving force for mass ow from nozzle to collector, but also to interfere
with fiber morphology, particularly diameter [14 ]. In addition, it is
essential that both FFF and collector conduct electricity. Actually, besides being a requisite for electrospinning, electrical conductivity of FFF
has been shown to play a role in fiber diameter and bead formation
[14 ,144]. The last major stage of the electrospinning apparatus, i e
fiber collection, is also essential for tailoring the properties of the produced ultrathin fibers. hile electrical conductivity is a requirement
also for the collecting surface, its shape [145] and mobility – static,
when absent, or mobile (usually rotating) – also determine the properties of the resulting materials, particularly diameter, bead formation,
degree of crystallinity, and the presence or absence of preferential molecular orientation. Rotation speeds ranging from 15 to 4
rpm
[145,14 ] have been reported. Table
summarizes FFF and most
important processing parameters.
erreira et al
Fig. . Electron micrographs of electrospun polymer-based ultrathin fibers: SEM of (a) PHB-based and (b) PHB hydroxyapatite fibers. (c) TEM of PBAT-based fibers
containing hydroxyapatite. SEM of (d) PLLA and (e) PLLA-BG fibers. (f) SEM of PBAT polypyrrole blends showing beads (red arrow). (For interpretation of the
references to color in this figure legend, the reader is referred to the web version of this article.)
Figures (a–b) and (c) were adapted from Ref. [15 ] and [149], respectively, with permission from Elsevier. Figures (d-e) were adapted from Ref. [1 ] with
permission from ACS Copyright © 11 American Chemical Society. Figure (f) was adapted from Ref. [15 ] with permission from The Royal Society of Chemistry.
Characterizing electrosp n polymer ltrathin
ers
Conventionally, the biomedical materials intended for bone tissue
engineering ought to feature a highly porous structure and a rough
surface to promote cell function [ , ]. Moreover, these materials
should present mechanical properties similar to the bone tissue in order
to assist in bone regeneration, as well as good performance in both in
vitro cell culture and in vivo implantation [15 ]. Several characterizations are usually carried out to assess the properties of electrospun
polymer ultrathin fibers. The most important ones are discussed below.
Morphological characterization
The morphology of electrospun ultrathin fibers has been studied
using scanning electron microscopy (SEM; Fig. a–b) and transmission
electron microscopy (TEM; Fig. c) [149,15 ,15 ,15 ,158]. Both techniques allow the observation of highly porous structure and the presence
of bioactive ceramics decorating the fibers (Fig. d–e). Moreover, such
microscopies also allow observing hydroxycarbonate apatite formation
after immersion in simulated body uid (SBF) [151]. These features are
important because they facilitate cell adhesion. Arslan et al. [159] reported that highly porous, rough PBAT-based electrospun ultrathin fibers facilitated the attachment and spread of osteoblastic cells (MC T E1). The micrographs (Fig. f) also show beads in the meshes, the former
being related to the viscoelasticity of the solution, charge density carried
by the jet, and the surface tension of the solution [1 ]. Deitzel et al.
[1 1] reported that the current of the electric field has also an in uence
on the formation of beads within the fibers. According to Fong et al.
[1 4], higher polymer concentration results in fewer beads.
Mechanical characterization
Advances in scaffold manufacturing are emerging and new scaffolds
mimic ing the dynamic, complex architecture of the extracellular matrix have been reported [1 ,1 4]. In this context, the mechanical
performance of the fibers is important to facilitate the connection of the
scaffold with the surrounding tissue [114]. The mechanical properties of
electrospun polymer ultrathin fibers have been studied theoretically
[1 5] and experimentally by stress-strain curves [151,1 ]. Several
authors have reported that the addition of bioactive ceramics can
improve the mechanical performance of electrospun polymer ultrathin
fibers [1 ]. Jeong et al. [1 8] prepared electrospun fibers with 5 and
1
HA and observed that the tensile strength increased from .
±
. 4 MPa to .
± .
MPa after addition of 1 wt HA. Similar
results were reported by Chen et al. [15 ], who observed an increase
from 1. ± . 5 MPa to .99 ± .5 MPa in the tensile strength at brea
after addition of 5 wt nHA into electrospun PHB fibers.
On the other hand, other groups have shown negligible or detrimental effects in overall mechanical properties of the electrospun
polymer ultrathin fibers with the addition of bioactive ceramics. For
instance, McCullen et al. [1 9] added β-tricalcium phosphate (TCP) into
PLA and observed that the tensile strength of the electrospun fibers
decreased from 84 ± 89 Pa to 5 ± 9 Pa (ultrathin fibers containing 5 wt TCP). Liverani et al. [151] observed a decrease from ± 1
MPa to . ± . 4 MPa in the oung’s modulus of nanofibers based on
PCL chitosan after adding
wt BG. Table summarizes some mechanical results of electrospun polymer ultrathin fibers containing
bioactive ceramics.
From the aforementioned contradictory results, it is clear that
Table 3
Mechanical properties of electrospun polymer ultrathin fibers containing bioactive ceramics.
Organic matrix
Inorganic filler
Content of filler (wt )
Tensile strength (MPa)
PLA
PLA
PLA
PLA
PCL
PCL chitosan
PHB
HA
TCP
HA
PLA- g-HA
Nanosilicate
BG
nHA
5–
5–
. 5, .5
–
1–1
.
.84
.5
.
4.5
5
1.
Neat polymer
oung’s modulus (MPa)
Composite
.15 – .
. 9– .44
.5 – . 5
1. – .
18–5
.99
Neat polymer
.419
.5
–
55
.1
5
Ref.
Composites
1.8–4.
4.5–8.5
–
–9
. 8– .84
.
[1 8]
[1 9]
[1 ]
[1 1]
[1 ]
[151]
[15 ]
erreira et al
Fig. . (a) In vitro cell culture showing mineralized nodule formation after 14 days of culture. (a.1) PCL, (a. ) PCL nHA (
). (b) Confocal laser scanning microscopy images of MG- osteoblast-li e cells cultured for 8 h on (b.1) PHB PCL and (b. ) PHB PCL BG scaffolds. (c) Schematic of in vitro cell adhesion. (d) SEM
micrographs of (d.1) PBAT and (d. ) PBAT PPy nHAp showing the MG- cells attached on the scaffolds surface (blue painted). Scale bar = 1 μm. (For interpretation of the references to color in this figure legend, the reader is referred to the web version of this article.)
(a) Adapted from ref. [ ] with permission from Elsevier. (b) Adapted from ref. [154] ACS Copyright © 1 American Chemical Society. (c) Based on the wor of
Ref. [11 ]. (d) Adapted from ref. [15 ] with permission from The Royal Society of Chemistry.
ceramic materials affect the mechanical properties of the nanofibers as
well as that there is an optimal concentration above which any further
increase in the content of bioactive ceramics will harm the mechanical
properties of the scaffolds [1 ]. For instance, u et al. [1 1] prepared
PLA-based electrospun fibers containing different concentrations of
PLA-grafted HA (ranging from to
wt ). The authors observed the
presence of an optimal point (1 wt ) where the tensile strength is
greater, and then decreases as the content increases. ang et al. [1 ]
also observed an optimal point of mechanical properties (5 wt nanosilicate) of the PCL-based fibrous scaffolds fabricated via electrospinning. Therefore, it is recommended to study the effect of the
bioactive ceramics in each system, because it can vary and modify the
mechanical properties of the final material.
The mechanical performance of the electrospun polymer ultrathin
fibers can also be controlled by tailoring the fiber orientation during the
preparation, as observed by Lee et al. [1 ]. Preferential alignment
shows highly anisotropic mechanical strength particularly in the alignment direction.
In vitro cytoto icity
Electrospun polymer ultrathin fibers have shown promising
morphology, structure, and mechanical properties, all being suitable for
tissue engineering. However, cell viability needs to be carefully studied
to avoid compromising cellular functions. Only non-cytotoxic ultrathin
fibers are suitable for biomedical use in accordance with ISO-1 99 -5
[1 4]. The most versatile and popular tests to assess in vitro cell
viability are MTT and live dead viability cytotoxicity assays [1 5,1 ].
These assays demonstrate if there is a significant statistical difference in
cell viability between the control group and the membranes. In the case
of MTT viability assay, the conversion of the water soluble MTT ( -(4,5dimethylthiazol- -yl)- ,5-diphenyltetrazolium bromide) compound to
an insoluble formazan is analyzed [1 ]. The level of cell viability can
vary. Thomas et al. [1 8] studied the viability cytotoxicity assays of
PLA by MTT and used ≤
of cell viability as the value from which a
cytotoxic effect is considered. Other authors [1 9] classified a material
as cytotoxic if 5
of cell viability is injured. In the case of live dead
viability cytotoxicity assays, the living and dead cells after contact with
the material are quantified and the results are compared with the
quantification results of control [14 ].
. Electros un olymer ultrat in
engineering
bers for bone tissue
In vitro cell c lt re
Cell culture has been used to evaluate the biological properties of
biomaterials in bone regeneration. The in vitro performance of materials
is usually studied by adhesion of different cell lines and their subsequent
differentiation into osteoblasts after some days. A number of authors
have reported better cellular response during in vitro cell culture of
electrospun polymer ultrathin fibers containing HA [18 ,181] and BG
[1 9,18 ,18 ]. Santana-Melo et al. [ ] reported that the osteoconductive and osteoinductive properties of nHA ( and 5 wt ) present
in PBAT-based ultrathin fibers improved the gene expression related to
osteogenesis of osteoblast-li e MG cells (Fig. a). Other authors [181]
observed that the addition of nHA ( , 5, and
wt ) increased the
cellular differentiation (rat mesenchymal stem cells — rMSCs), indicating that the addition of nHA improves the performance of electrospun
ultrathin fibers during differentiation.
Shamsi et al. [14 ] reported that the bioactive ions from the glass
present in PLLA BG ultrathin fibers improved cell attachment, growth,
proliferation, and gene expression. Other authors [184] prepared PLA
BG ultrathin fibers and showed that osteoblastic cells (MC T -E1)
adhered better on the composite ultrathin fibers. Similar results were
observed by Ding et al. [154], who reported that ultrathin fibers of PHB
PCL containing BG are suitable for MGcell adhesion and also
improve d cell viability (Fig. b). These promising results are attributed
to the large specific surface areas of electrospun polymer ultrathin fibers
containing bioactive ceramics, which can be used as binding sites to cell
membrane receptors [185]. This cell adhesion is important because it
regulates the differentiation and migration of cells [18 ].
In vitro cell adhesion occurs in three phases (Fig. c): attachment,
attening, and spreading [11 ]. In the first phase, electrostatic interaction and slight attening are observed [18 ]. During the second
phase, the cell attens and spreads on the substrate. Maximum spread
area of cell is observed during the third phase, where the adhesion is
stronger between cells and ECM due to electrostatic interactions [188].
In vivo cell adhesion is a dynamic process with “doc ing” and “loc ing”
phases [11 ]. More details on cell adhesion for biomedical applications
can be found elsewhere [11 ,189–19 ].
Several factors can affect cell behavior and consequently its osteogenic differentiation, including electrical conductivity [15 ] and
erreira et al
Table 4
Studies on the incorporation of active components into electrospun ultrathin polymer fibers that effectively increase regulation of osteogenic genes.
Polymer
Active components
Fabrication method
Effects on gene expression
Ref.
PLLA
PCL
Lactoferrin
Electrospinning +
immobilization of Lactoferrin
[
PLA
PEG
PLGA
Dexamethasone (DE ) and bone
morphogenetic protein- (BMP- )
BMP- n HA
Coaxial electrospinning
technology
Electrospinning
PCL
Micro-RNAs miR
Dual power electrospinning
PLGA
PCL
Octacalcium phosphate (OCP)
Electrospinning
PLACL*
Blending and spraying
PLGA
Sil fibroin (SF) and hydroxyapatite
(HA)
HA
PCL
CA-HAmicroparticles
Up regulation of IGF-1R, RANKL, Runx , OPN and OCN inducing osteogenic
differentiation of human adipose-derived stem cells (ADSCs) cultured on the
lactoferrin-coated membranes
Dual-bioactive molecule-loaded fibrous meshes exhibited a significantly higher
expression level of ALP, OPN, and OCN in bone mesenchymal stem cells (BMSCs)
The expression of OCN, Runx , and ALP genes was higher on BMSCs seeded onto
random fiber regions. Sox9 expression (mar er of chondrogenic differentiation)
higher cells seeded on neatly aligned regions
iPSCs seeded on the scaffold showed higher expression of Runx , OCN, ALP, and
Osteoconectin leading to osteogenic differentiation and matrix mineralization
Rat mesenchymal stem cells (MSCs) seeded onto the OCP-loaded membranes for up
to 14 days expressed higher levels of Runx , ALP, COL1a1, OPN, OCN, and BMP
leading to osteogenic differentiation and matrix mineralization
Enhanced ALP and OCN expression resulting in greater osteogenic differentiation
of mesenchymal stem cells (MSCs) on PLACL SF HA(s) compared to all other NFS
PLGA HA Col scaffolds stimulated greater up regulation of ALP, OCN, and Runx
in rat MSCs promoting osteogenic differentiation
The composites with aligned fibers and
wt Ca promoted higher expression of
ALP enhancing osteogenic differentiation
*
and miR-1
Electrospinning, biomimetic
process, and adsorption
Electrospinning
4]
[194]
[
]
[195]
[19 ]
[19 ]
[198]
[199]
PLACL — poly(L-lactic acid)-co-poly( -caprolactone).
morphology of fibers [159,19 ]. Bashur et al. [19 ] reported that the
alignment and orientation of the fibers affect the cell morphology, while
the expression of mar ers is more sensitive to their diameter and length.
Similar observations were made by Arslan et al. [159], who reported
that the roughness and higher surface area of PBAT-ultrathin fibers
facilitate the attachment and spread of MC T -E1 pre-osteoblastic cell
line, and the pore size (<1 μm) of the ultrathin fibers hampers cell
infiltration. In order to increase the pore size and porosity, Chen et al.
[15 ] prepared PHB-based ultrathin fibers containing nHA by electrospinning and then the thin layers of nHA PHB were folded three times to
obtain scaffold with eight layers. They reported that the laminated nHA
PHB scaffolds showed greater cell-loading capacity in addition to
improved cellular functions. Castro et al. [15 ] observed that PBAT
polypyrrole scaffolds presented higher values of al aline phosphatase
besides an increase in electrical conductivity, when compared to pure
PBAT, suggesting improvement in osteoblast differentiation (Fig. d).
Recent investigations have demonstrated that the incorporation of
active components into electrospun ultrathin polymer fibers significantly increases the up regulation of osteogenic genes, such as Al aline
Phosphatase (ALP), Runt-related transcription factor
(Runx ),
Collagen (Col1a1), Osteocalcin (OCN), and Osteopontin (OPN) (Table 4)
[194– 1]. The ALP gene expression level is widely used as a mar er to
identify cells in the early stage of osteogenic differentiation. Runx , also
nown as core-binding factor subunit alpha-1 (CBFα1), encodes a transcription factor that regulates many other genes related to osteogenesis,
thus, directly lin ed to bone formation. Col1a1 is a gene that contains
instructions for ma ing part of type-I collagen, which is the most
abundant protein of the bone extracellular matrix. OCN and OPN are
used as late mar ers of osteogenic differentiation and the proteins
encoded by these genes are related to the process of matrix mineralization [
].
Loading electrospun fibers with active agents seems to enhance
osteogenic gene expression by physical or biological routes depending
on the doping agent. Deng et al. [
] have demonstrated that incorporating polydopamine nanoparticles into fibrous membranes significantly increases its hydrophilicity. This change led to a higher
adsorption of adhesive proteins onto the membrane’s surface ultimately
easing cell attachment, spread, and differentiation. A similar hypothesis
is proposed for HA-containing scaffolds. HA-coated PGLA presented
enhanced hydrophilicity and protein adsorption when compared to pure
PGLA scaffolds, resulting in greater expression of ALP, OCN, and Runx
[
].
Immobilizing proteins such as BMP- , Dexamethasone, collagen,
lactoferrin, or micro RNAs (miR
and miR-1 ) in electrospun
polymeric meshes was shown to promote strong up regulation of osteogenic genes (Table 4). Each of these molecules uses different intracellular pathways to transmit the message to the cell to carry out the gene
expression and subsequent translation of osteogenic proteins. In a recent
investigation, Lee et al. [ 4] have successfully immobilized lactoferrin
in PLLA and PCL nanofibrous membranes. After culturing human
adipose-derived stem cells (ADSCs) onto these membranes they
observed a higher expression of insulin-li e growth factor receptor 1
(IGF1R), RANKL, Runx , OPN, and OCN with pronounced osteogenic
differentiation. In spite the exact mechanism by which lactoferrin
stimulates osteogenic differentiation needs elucidation, some evidences
suggest that it could activate the cAMP signaling pathway that induces
OCN expression in mouse osteoblasts [ 5]. Ta en together these results
demonstrate that the development of multi-functional scaffolds that
mimic both the physical and biological environment of the bone extracellular matrix may result in enhanced osteogenic differentiation.
In vivo implantation
D electrospun-based scaffolds have demonstrated potential for in
vivo implantation due to their structure and mechanical properties.
However, these scaffolds have low bioactivity, which is a property often
associated with bone growth. As previously mentioned, a current
strategy in bone tissue engineering is the addition of bioactive ceramic
to prepare bioactive polymer ultrathin fibers, especially HA and BG. The
bonding of material with the surrounding bone tissue is therefore
enhanced. Several authors have reported promising in vivo implantation
results of D electrospun-based scaffolds containing HA and BG. Chen
et al. [15 ] prepared PHB-based scaffolds containing nHA (5 w v) and
then MSCs were seeded onto the scaffolds to fabricate bone grafts. They
observed that both PHB and PHB nHA scaffolds were encapsulated with
a thin layer of soft tissue (Fig. 8a) after
months of implantation.
Moreover, the addition of nHA improved collagen deposition (blac
arrow in Fig. 8a.5 and a. ) and blood vessel density at the implant site.
Silva et al. [148] studied the in vivo implantation of PBAT ultrathin
fibers containing different concentrations of nHA (1, , , 4, 5, and wt
) on tibia bone defects. They observed that PBAT filled with wt
nHA improved bone volume, force, and stiffness when compared to
control group. Other authors [181], also studying tibia bone defects,
prepared PCL-based ultrathin fibers with different amounts of nHA ( , 5,
and
wt ) observed a linear closure (arrows in Fig. 8b) only for the
group treated by PCL nHA (5 ). Santana-Melo et al. [ ] observed that
independently of concentration ( and 5 wt nHA), nHA-loaded PBAT
ultrathin fibers showed improved bone repair when compared to neat
erreira et al
Fig. . (a.1–a.4) Representative images showing the scaffolds encapsulated with a thin layer of soft tissue. (a.5 and a. ) Histological staining with Masson’s trichrome staining (MTS); blac arrows indicate vascular structures. (b) Histological Section ( ×) of bone repair after 4 wee s of implantation showing linear closure
(arrows) only for group treated by PCL nHA (5 ). (b.1) PCL, (b. ) PCL nHA (5 ) and (b. ) PCL nHA (
) scaffolds. (c.1 and c. ) Representative images of subcutaneous implant. Histological analysis of subcutaneous tissue of the rats after
days of treatment (c. and c.4) neat membrane and (c.5 and c. ) membrane
containing HA. (d.1) Critical-size defect created in rat calvaria, (a. ) defect without and (a. ) with implanted scaffold after 8 wee s. Multislice spiral-computed
tomography images after 8 wee s of study: (d.4) control, (d.5) PLLA, (d. ) PLLA HA, (d. ) PLLA BG, (d.8) PLLA BG HA. Optical micrographs of the defects:
(d.9 and d.14) control, (d.1 and d.15) PLLA, (d.11 and d.1 ) PLLA HA, (d.1 and d.1 ) PLLA BG, (d.1 and d.18) PLLA BG HA. Magnification of (d.9 to d.1 ) 1
and (d.14 to d.18).
(a) Adapted from ref. [15 ] with permission from Elsevier. (b) Adapted from ref. [181] with permission from Springer. (c) Adapted from ref. [149] with permission
from Elsevier. (d) Adapted from ref. [1 ] ACS Copyright © 11 American Chemical Society.
polymer. This improvement after addition of nHA was observed by the
same authors in previous studies when PDLLA was used as matrix [
].
Importantly, Ribeiro Neto et al. [149] studying the in vivo performance
of bilayer scaffold prepared by PBAT-based film HA and nanofiber HA
observed a mild in ammatory response after implantation (Fig. 8c).
BG has also been used during the preparation of scaffolds and
improved performance has been observed. Ardeshirylajimi et al. [15 ]
reported that the BG in polyethersulfone (PES) ultrathin fibers enhanced
the osteogenic mar ers of MG- cells and the bone formation in vivo.
Dinarvand et al. [1 ] reported a different approach for electrospun
polymer ultrathin fibers containing bioactive ceramics, preparing PLLA
scaffolds coated with both HA and BG. The authors observed improvement in bone reconstruction for the animals treated with HA- and BGcoated PLLA scaffolds. However, the best results were observed when
the scaffolds were simultaneously coated with HA and BG (Fig. 8d.8).
Moreover, different from results observed by Ribeiro Neto et al. [149],
they reported that the scaffolds do not induce in ammation at the
implanted site. Table 5 summarizes the in vitro and in vivo results of
electrospun polymer ultrathin fibers containing bioactive ceramics for
bone tissue engineering.
A number of synthetic polymers (bioresorbable or not) have been
used to prepare the ultrathin fibers, although clinical demand for implants has focused a great deal on bioresorbable polymers. The preparation of electrospun polymer ultrathin fibers using these polymers is
still a challenge due to the difficulty in homogenizing the inorganic
phase (hydrophilic) in the organic phase (hydrophobic). In this context,
several authors have exploited the use of smaller bioactive ceramics
(reaching the nanoscale), surfactants and surface modified bioactive
ceramics as an attempt to overcome the formation of inorganic nanoparticle beads during electrospinning [155, 9]. Several authors have
shown excellent nanofiber performances in both in vitro cell culture and
in vivo implantation. However, performance , under clinical conditions
erreira et al
Table 5
Electrospun polymer ultrathin fibers containing bioactive ceramics for bone
tissue engineering.
Polymer
matrix
Amount of
bioactive
ceramics
In vitro-tested cells
In vivo protocol
Ref.
PCL
MSCs
PHB
5 wt
MSCs
PBAT
and 5 wt
nHA
Bioactive
glass
nanofiber
nHA*
MG
Tibia bone
defects
Tibia bone
defects
Dorsal
subcutaneous
poc ets
Tibia bone
defects
–
[181]
PBAT
, 5, and
wt nHA
1– wt nHA
–
[15 ]
5, 1 , and
BG
and
wt
nHA**
BG-coated
MG-
–
[154]
hASC
[149]
BG coated
with PDLLA
layer
HA- and BGcoated
BG
BG nanofiber
BG nanofiber
Chondrocyte cells
Sub-cutaneous
implant
Calvarial bone
defect
–
[
]
Calvarial bone
defect
–
–
–
[1
]
PLLA
PBAT
PPy
PHB PCL
PBAT
PES
PDLLA
PLLA
PLA
PLA
PCL
*
**
nHA
MG-
Human bone marrow
mesenchymal stem
cells
MG-
MG-
–
MC T -E1
MC T -E1
MC T -E1
[148]
[15 ]
[
]
[14 ]
[15 ]
[184]
[18 ]
[ 8]
nHA crystals were electrodeposited on PBAT PPy scaffolds.
Bilayer composite structures.
has yet to be investigated in order to incorporate this technology into
clinical practice.
.
emar s and ers ectives
Polymer-based ultrathin fibers are promising materials for biomedical applications and, as such, have received significant attention in
recent years. The combination of these ultrathin fibers loaded with
bioactive ceramics is a great route to prepare scaffolds with suitable
properties for bone regeneration. In practice, a synergistic effect occurs,
i e the bioactive ceramics improve the adhesion, differentiation, and
calcification between the bone matrix and the ultrathin fibers, while the
brittleness of bioactive ceramics (major limitation for their use as biomaterials) is overcome by the mechanical exibility of the polymer fibers. Novel technologies to fabricate polymer inorganic composite and
hybrid ultrathin fibers have been reported in the literature and have
opened the path for the preparation of ultrathin fibers with unique
performance. Electrospun polymer ultrathin fibers containing bioactive
ceramics have potential to mimic native ECM, promoting cell functions
of a wide variety of cell types. Nevertheless, further in vitro and in vivo
studies are needed to better understand the synergy between bone
regeneration and engineering. Overall, the vision for the future in the
field of polymer ultrathin fibers containing bioactive ceramics for bone
tissue engineering relies in developing stronger collaborations among
universities, companies, and hospitals to further enhance nowledge
and potential to achieve the desired results while also evaluating scaledup technologies to enable efficient technology transfer and benefit to
society.
Further spinning techniques have emerged in the context of polymerbased ultrathin fibers for tissue engineering applications [ 1 , 11].
These are mostly derived from the need to increase yield, one of the
major drawbac s of the electrospinning process. Solution blow spinning
(SB-Spinning), for instance, is remar ably similar to electrospinning
except that the force that draws FFF into a jet towards the collector
comes from a pressurized gas vs the electric field used in electrospinning. In other words, the need for applying high electrical fields is
eliminated in SB-Spinning, allowing ultrathin fibers to be in sit deposited onto a broader range of collectors – including non-conductive surfaces such as biological living tissues – with a much higher yield
[ 1 , 1 ]. Other example is the process called Forcespinning®, which
uses centrifugal force rather than electrostatic force as in the electrospinning process [ 14, 15]. Surface modification, including cold
plasma treatment, is another example of a technique that has emerged in
the context of electrospun nanofibers used in biomedical applications.
Positive effect on nanofibers properties (biocompatibility, immobilization adsorption of molecules of interest, and surface grafting crosslin ing) have been observed due to such treatment [ 11].
It is worth pointing out that each spinning techniques has advantages
and drawbac s, and contemporary combinations among these exciting
techniques have led to improved fiber-forming protocols, some of them
focused on bone tissue regeneration. Overall, the combination of
different spinning principles and unit operations into innovative fiberforming techniques is virtually endless and the progress obtained by
the scientific community in this direction has multiplied the possibilities
of such promising systems for a wide range of applications. It is now
envisaged that these engineered spinning processes can be also translated into bioactive ceramic-containing polymer-based ultrathin fibers,
further spreading their potential applications for bone regeneration
purposes.
eclaration of com eting interest
The author reports no con icts of interest in this wor .
c no ledgements
The authors ac nowledge Funda ão de Amparo à Pesquisa do Estado
de São Paulo (FAPESP - Grant 1
9588-9; Ph.D. fellowship of F.V.F.)
and Conselho Nacional de Desenvolvimento Cient fico e Tecnológico
(CNPq – Grant 1 88
- and 4 4 8
18-5 to AOL) for financial support. e are than ful to Prof. A. R. Boccaccini and Dr. L. Liverani
(University of Erlangen-Nuremberg, Germany) for expert comments on
the manuscript. Figs. , 4, 5 and c were created with BioRender.com.
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