Bioactive Materials 4 (2019) 271–292
Contents lists available at ScienceDirect
Bioactive Materials
journal homepage: http://www.keaipublishing.com/biomat
Recent advances in biomaterials for 3D scaffolds: A review
a,∗
b,c,d
Maria P. Nikolova , Murthy S. Chavali
T
a
Department of Material Science and Technology, University of Ruse “A. Kanchev”, 8 Studentska Str., 7000, Ruse, Bulgaria
Shree Velagapudi Ramakrishna Memorial College (PG Studies, Autonomous), Nagaram, 522268, Guntur District, India
PG Department of Chemistry, Dharma Appa Rao College, Nuzvid, 521201, Krishna District, India
d
MCETRC, Tenali, 522201, Guntur District, Andhra Pradesh, India
b
c
A R TICL E INFO
A BSTR A CT
Keywords:
Bioactive scaffolds
Bone tissue engineering
Polymeric biomaterials
Bioceramics
Bioprinting
Considering the advantages and disadvantages of biomaterials used for the production of 3D scaffolds for tissue
engineering, new strategies for designing advanced functional biomimetic structures have been reviewed. We
offer a comprehensive summary of recent trends in development of single- (metal, ceramics and polymers),
composite-type and cell-laden scaffolds that in addition to mechanical support, promote simultaneous tissue
growth, and deliver different molecules (growth factors, cytokines, bioactive ions, genes, drugs, antibiotics, etc.)
or cells with therapeutic or facilitating regeneration effect. The paper briefly focuses on divers 3D bioprinting
constructs and the challenges they face. Based on their application in hard and soft tissue engineering, in vitro
and in vivo effects triggered by the structural and biological functionalized biomaterials are underlined. The
authors discuss the future outlook for the development of bioactive scaffolds that could pave the way for their
successful imposing in clinical therapy.
1. Introduction
Implantable 3D scaffolds are used for restoration and reconstruction
of different anatomical defects of complex organs and functional tissues. The scaffolds provide a template for the reconstruction of defects
while promoting cell attachment, proliferation, extracellular matrix
generation, restoration of vessels, nerves, muscles, bones, etc. Scaffolds
are three-dimensional (3D) porous, fibrous or permeable biomaterials
intended to permit transport of body liquids and gases, promote cell
interaction, viability and extracellular matrix (ECM) deposition with
minimum inflammation and toxicity while bio-degrading at a certain
controlled rate. The artificial material substituted for tissue grafts is
called alloplastic. The alloplastic bioactive scaffolds ensure not only the
mechanical support of the tissue but also serve as a delivery vehicle for
bioactive molecules (cytokines, inhibitors, drugs, antibiotics, etc.) and
templates for attaching genetically transduced cells establishing new
centres for tissue regeneration and morphogenesis. Simultaneously, 3D
scaffolds can be used as tissue models replicating the structural complexity of the living tissues. For that reason, not only the biomaterial
used, but also the macro-, micro-, and nano-architecture of the scaffolds
are of prime importance.
Based on their chemical composition, biomaterials used for 3D
scaffolds are classified into metals, ceramics and glass-ceramics, natural
and synthetic polymers, and composites. Recently, the focus is put towards biodegradable biomaterials that do not need to be explanted
from the organism. Table 1 summarizes the main classes of biomaterials
used in 3D scaffold production together with their common application
and synthesis methods. Each biomaterial has specific chemical, physical, and mechanical properties, the ability for processing and control
of 3D shapes and geometry. The selection of production technique depends on the specific requirements for the scaffold, material of interest,
and machine limitation [1]. The integration of computer-aided design
(CAD) software and rapid prototyping proposes the ability to produce
objects with macro- (overall size and shape), micro- (pore size, shape,
interconnection, and distribution) and sometimes nanoarchitecture
(nano-roughness, topology, etc.) control of highly complex biomedical
devices. Using patient data, the design of the scaffold could be individualized by preparing special 3D model with certain porosity or
structures for vasculature that is compatible with multiple biomaterials
and cells. The use of 3D printing has revolutionized the development of
regenerative medicine and pharmaceutical field.
Recently, much research has been done to develop a variety of novel
biomaterials and composites with enhanced cell viability, cell proliferation, and printability. For additional modification, biomimicry
approach arranging different cellular components (growth factors,
hormones, ECM proteins, etc.) to mimic living tissue is used to enhance
Peer review under responsibility of KeAi Communications Co., Ltd.
∗
Corresponding author. Depatrment of Material Science and Technology, University of Ruse "Angel Kanchev", Bulgaria.
E-mail addresses: mpnikolova@uni-ruse.bg (M.P. Nikolova), ChavaliM@gmail.com (M.S. Chavali).
https://doi.org/10.1016/j.bioactmat.2019.10.005
Received 19 August 2019; Received in revised form 7 October 2019; Accepted 15 October 2019
Available online 25 October 2019
2452-199X/ This is an open access article under the CC BY-NC-ND license (http://creativecommons.org/licenses/BY-NC-ND/4.0/).
Bioactive Materials 4 (2019) 271–292
M.P. Nikolova and M.S. Chavali
MRI
magnetic resonance imaging
MSC
mesenchymal stem cells
MWCNTs multi-wall carbon nanotubes
NHEKs normal human epidermal keratinocytes
oMSCs
ovine mesenchymal stem cell
PAA
poly(acrylic acid)
PCL
ε-poly(caprolactone)
PEG
poly(ethylene glycol)
PEI
polyethyleneimine
PEO
poly(ethylene oxide)
PES
polyethersulphone
PET
polyethylene terephthalate
PGA
poly(glycolic acid)
PGF
platelet-derived growth factor
PHB
polyhydroxybutyrate
PHBV
poly(3-hydroxybutyrate-co-3-hydroxyvalerate)
PHEMA polyhydroxy ethyl methacrylate
PHPMA N-(2-Hydroxypropyl)methacrylamide
PLC
poly(L-lactide-co-ε-caprolactone)
PLGA
poly(lactic-co-glycolic) acid
PLA
poly(lactic acid)
PLLA
poly(L-lactic acid)
P(LLA-CL) poly(L-lactic acid-co–caprolactone)
PMMA
poly(methyl methacrylate)
PP
polypropylene
PU
polyurethane
PVA
polyvinyl alcohol
PVAc
polyvinyl acetate
rBMSC
rat bone mesenchymal stem cells
rCCs
rabbit corneal cells
rGO
reduced graphene oxide
rhBMP-2 recombinant human bone morphogenic protein-2
RNA
ribonucleic acid
SDSCs
synovium-derived stem cells
SH-SY5Y human-derived cells used as models for neuronal function
and differentiation
SiHAp
silicate containing hydroxyapatite
SLA
stereolithography
SLS
selective laser sintering
TGF-β1 transforming growth factor-β1
TPU
thermoplastic polyurethane
VEGF
vascular endothelial growth factor
β-TCP
β-tricalcium phosphate
Abbreviations
2D
3D
3T3
ALP
ADMSCs
BG
bFGF
BMP-2
BMSCs
BSA
CPC
CSH
CT
dECM
DMSO
DNA
EBM
ECM
EGF
EPCs
FHAp
FDM
GF
GO
HA
HAp
hADSCs
hAVICs
hBMSCs
hDFCs
hEKCs
hFOBCs
hNDFs
hTMSCs
two-dimensional
three-dimensional
fibroblasts cells
alkaline phosphatase
adipose tissue-derived mesenchymal stem cells
bioactive glass
basic fibroblast growth factor
bone morphogenic protein-2
bone marrow stromal cells
bovine serum albumin
calcium phosphate cement
calcium sulphate hydrate
computed tomography
decellularized extracellular matrix
dimethyl sulfoxide
deoxyribonucleic acid
electron-beam melting
extracellular matrix
epidermal growth factor
endothelial progenitor cells
fluorhydroxyaptite
fused deposition modelling
growth factor
graphene oxide
hyaluronic acid
hydroxyapatite
human adipose-derived stem cells
human aortic valvular interstitial cells
human bone marrow-derived stem cells
human dental follicle cells
human embryonic kidney cells
human fetal osteoblast cells
human neonatal dermal fibroblasts
human inferior turbinate-tissue derived mesenchymal
stromal cells
hUCMSC human umbilical cord mesenchymal stem cells
hUVECs human umbilical vein endothelial cells
hUVSMCs human umbilical vein smooth muscle cells
LBM
laser beam melting
MBG
mesoporous bioactive glass
MC3T3-E1 osteoblast precursor cell line
MG-63
human osteosarcoma cell line
biocompatibility, etc.) requirements as well as considerations concerning sterilization and commercial feasibility. To improve the
bioactivity and functionality of 3D scaffolds acting as synthetic frameworks or matrices, the shape, size, strength, porosity, and degradation
rate are readily controlled. The design of these regeneration templates
has evolved over the past years. To repair the damaged tissue, the
scaffold should be designed and fabricated in a manner resembling the
anatomical structure and mimicking the function and biomechanics of
the original tissue. The 3D scaffold should temporarily withstand the
external loads and stresses caused by the formation of the new tissue
while preserving mechanical properties close to that of the surrounding
tissue. It was demonstrated that the tissue-specific mechanical characteristics, in particular, stiffness, could control the differentiation of
MSCs [3]. Simultaneously, the scaffold designs such as sponges, meshes, foams, etc., are able the control biodegradation as a key factor in
tissue engineering. The degradation of biomaterials could be surface or
bulk. In contrast to bulk degradation that breaks the internal structure
of the material, the surface degradation maintains the bulk structure.
The rate of degradation should match the tissue growth without
cell signalling or ECM formation [2]. Moreover, biomaterial scaffolds
are used for delivering therapeutic agents like proteins, growth factors,
drugs, etc. and the anchorage of these substances to the scaffold is of
high importance for loading. As biomaterial-cell interactions are key to
cell viability, proliferation and differentiation, characteristics of biomaterials such as surface chemistry, charge, roughness, reactivity, hydrophilicity, and rigidity need to be considered. With continuous research and progress in biomaterials used in medical implants, the goal
of this review is to discuss recently developed implantable scaffold
materials for different tissues and highlight the advances in in vitro and
in vivo biological regeneration.
2. 3D scaffold requirements
A large number of scaffolds with various macro- and microarchitectures from different biomaterials have been reported in the
literature. The design of a scaffold includes mechanical (stiffness,
elastic modulus, etc.), physicochemical (surface chemistry, porosity,
biodegradation, etc.), and biological (cell adhesion, vascularization,
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Table 1
Biomaterials for 3D scaffolds production together with their common application and fabrication methods.
Class biomaterial
Biomaterial for 3D
scaffolds
Application
CERAMICS
(HA, β-TCP, α-TCP, ZrO2, TiO2, porous bioglass, calcium silicate,
calcium sulphate, etc.)
Hard tissue replacement
Orthodontic application
POLYMERS
Natural
Connective and hard tissue
application
Decellularized living
tissues/organs
Drug delivery
Hard and soft tissue on
applicants
Gene therapy
Proteins (silk, collagen, gelatin, fibrinogen, actin,
keratin)
Polysaccharides (alginate, chitosan, cellulose,
dextran, chitin, glycosaminoglycan, hyaluronic acid,
agarose)
Polynucleotides (DNA, RNA)
Synthetic Degradable (polyesters, polyorthoesters, polylactones,
polycarbonates, polyanhydrides, polyphosphazenes,
etc.)
Non-degradable (PE, PTFE, PMA, PAA, PU,
polyether, polysiloxanes, etc.)
Drug-delivery systems
Implants
Orthopaedic implants
METALS &
ALLOYS
(Co–Cr, Ti, Ti–6Al–4V, stainless steel etc.)
Orthopaedic and dental
application
Artificial hearing
COMPOSITES
Blends of polymers and ceramics/metals
Orthopaedic and dental
application
separation of toxic byproducts. The degradation of a biomaterial could
be achieved by physical, chemical, biological or combined processes
influencing the biocompatibility of the 3D scaffold. For example, incorporating different biodegradable components in the construct triggers hydrolytic degradation while processes such as enzymatic digestion and cell-driven degradation biologically change the implant
material. When the application of a scaffold does not require a complete
degradation (for example in articular cartilage repair) permanent (nondegradable) or semi-permanent scaffolds could be used. When implanted in body right, toxic, immunological or foreign body responses
should not occur which prove the scaffold biocompatible. The surface
properties of a scaffold should also be designed in such a way that to
facilitate cell attachment, homogeneous distribution, proliferation and
cell-to-cell contacts. The scaffold geometry should maintain the porous
or fibrous design and provide high surface-to-volume ratio for cell attachment and tissue development. Nanostructured surfaces demonstrate high surface energy as opposed to polished materials that result
in enhanced hydrophilicity and, therefore, improved adhesion of proteins and cell attachment. For metal and ceramic scaffolds, the smaller
grain size not only increases the mechanical strength but was found to
be more favourable in terms of attachment and proliferation of osteogenic cells [4]. Therefore, the scaffold with its topography and mechanical features controls cellular behaviour. When seeded in 3D
scaffolds, cells need to be urged to regain typical in vivo morphology.
Fabrication
-
Binder jetting/inkjet printing
Extrusion
Stereolithography
SLS
Laser aid gelling
FDM
Polymer sponge replica
Salt leaching
Dual-phase leaching
Gel casting
Solvent casting
Inkjet printing
Particle aggregation
Micro moulding
Photolithography
Emulsification
Electrospinning
Cryo-gelation
Sol-gel
SLA
SLS/SLM
EBM
FDM
Polyjet
Electrospinning
Phase separation
Freeze drying
Gas foaming
Inverse opal hydrogelation
Self-assembly
SLA
SLM
SLS
EBM
Powder metallurgy
Vacuum foaming
Directional solidification
Textile base fabrication
Laminated object manufacturing (LOM)
FDM
SLA
Freeze-drying method
The process of regeneration also requires the development of interconnected neurovascular networks between the mature and surrounding tissue. On one hand, the scaffold design should make allowance for vascular remodelling as tissue mature so that nutrients, oxygen
and other soluble factors could reach all embedded cells while the
metabolic wastes are constantly removed. On the other hand, nerve
fibres are spatially closely associated with cells that express receptors
for neuropeptides and should be simultaneously developed with the
new tissue to regulate homeostasis. Usually, the distribution of peripheral nerves and blood vessels follows each other in human body
development because they are anatomically coupled and influence the
growth and development of each other [5]. Since it is still hard to
regulate multi-tissue types development, autologous neurovascular
bundles integrated by microsurgery during scaffold implantation is a
potential concept [6] for improving scaffold performance.
To support and accelerate the endogenous healing process, especially in extensive or irreversible damages, different strategies for administration of stem cells (after in vitro expanded) alone or in combinations with natural or synthetic scaffolds are proposed. Stem cells from
different sources (bone marrow, adipose, muscle tissue, lung, umbilical
cord, etc.) are usually used as therapeutic rely on because of their
ability to maintain homeostasis in healthy tissues and differentiate
when activated under a reparative response or disease. When using
tissue-specific stem cells, they are able to regenerate the tissue from
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[21]. These nanofibres mimic the structure and properties of the extracellular matrix and because of their high aspect ratio, porosity, and
surface-to-volume ratio, they induce greater cellular attachment than
microfibres [22]. The nanostructure may help the rapid diffusion of
encapsulated substances and cell infiltration [23]. When cultured on
smaller fibres (283 nm) of polyethersulphone (PES) with laminin
(protein of ECM, a major component of the basal lamina), rat neural
stem cells showed 40% increase in oligodendrocyte differentiation and
20% increase on larger fibres (749 nm) [24]. Commonly used techniques for fabrication of nanofibres are electrospinning, self-assembly,
phase separation, and solid free-form fabrication. In electrospinning, by
applying electrical field, the ejected material from the syringe pumps
forms fibres from nanometers to several micrometres in size producing
fibrous scaffolds. During electrospinning high voltage is imposed between the spinneret and the ground collector to thinner the filaments.
In order to have stable fibres, polymers with enhanced chain entanglement such as PCL, PLLA, PLGA, and polymer-containing composites, are used. The method has high loading and encapsulation capacity of small molecules and drugs [25]. Phase separation technique
separates phases through cooling or non-solvent exchange of heterogeneous fibre structures with little control over diameter and orientation of fibres whereas the self-assembly process autonomously organizes components into nanofibre structures. The phase separation
method suffers from drawbacks such as limited material selection and
inadequate resolution while self-assembly requires careful molecular
design because of the complex mechanism of synthesis. Recently, Yao
et al. proposed the production of scaffolds using thermally-induced selfagglomeration of short individual electrospun nanofibres and tiny nanofibre pieces in ethanol/water/gelatin (4/2/1) solution followed by
freeze-drying [26]. The hierarchical structure consisted of random
overlaid PLC/PLA fibres with a diameter ranging from 150 nm to 2 μm,
pores from sub-micron to around 300 μm, and about 96% porosity. It
was also found that the aligned fibres are able to direct the tissue
growth [27] while random fibres demonstrate improved stiffness in all
directions [28]. Yin et al. discovered that aligned PLLA scaffolds induced endogenic differentiation in MSCs while the cells cultured on
randomly orientated fibres displayed enhanced osteogenic differentiation [29]. These facts provide inside into development of fibrous scaffold design with smart functionalization in interfacial tissue engineering. For that reason Cai et al. developed dual-layer aligned
random nanofibrous scaffolds of electrospun silk fibroin-P (LLA-CL)
that effectively augmented tendon-to-bone integration and improved
which they are isolated. After the injury, a cascade of biological events
such as stem cell migration, chemokine and growth factor secretion
occurs to repair the destroyed tissue. These processes could be mimicked, stimulated and controlled by adding a variety of bioactive
moieties such as growth factors, peptides, genes, aptamers (single
strength oligonucleotides), antibodies, drugs, or even ECM. These
substances are chemically or physically (by electrostatic forces, hydrophobic interactions, and hydrogen bonds) decorated to the scaffold.
In that way, the engineered complex scaffolds mimic natural signalling
and repair events and generate suitable microenvironment for adhesion, proliferation, differentiation of stem cells that regenerate the
tissue.
3. Types of 3D scaffolds based on their geometry
Scaffolds of various biomaterials cover a wide range of applications
(Fig. 1). The classification of these complex constructs could be based
on their geometry or material source.
When the geometry is used to categorize 3D scaffolds, the available
biological constructs are:
3.1. Porous scaffolds
Sponge or foam porous scaffolds usually contain interconnected
pore structure (orientated or random) that is highly useful in bone regrowth, vascularization and ECM deposition. Because of the high
physical surface, such scaffolds provide improved gas and nutrients
transport through the channel network. However, increased pore interconnectivity is required for peripheral nerve and blood vessels
growth without scarifying the mechanical properties of the scaffold.
Otherwise, the cell in-growth and flow of nutrients will be prevented. If
the pores are too small, the cellular penetration, ECM deposition, and
neovascularization could be prevented. Ideally, the interconnected
porous structure should consist of 90% porosity [7] which could substantially influence the resulting mechanical properties. However, the
ideal size of pores for specific cells and tissues varies substantially [8].
For example, pores with 200–400 μm size were found to be effective for
bone tissue formation [9], while 50–200 μm pores were suitable for
smooth muscle cell growth [10]. Scaffolds with pore sizes between 1044 and 44–75 μm enable accommodation only of fibrous tissue [11]. As
a rule, pores with sizes greater than 100 μm enable tissue growth and
vascularization, whereas micro- (less than 2 nm) and mesopores
(2–50 nm) promote cell adhesion and resorbability at controllable rates
[12]. The pores that are excessively large in size (more than 400 μm)
decrease the cell-to-cell contact ratio since the cells experience twodimensional (2D) growth pattern on the substrate rather than a 3D
organization [6]. The special properties of porous scaffolds could be
obtained by using methods capable of applying control over pore sizes.
Except for the conventional freeze-drying and solvent casting/particulate leaching techniques, fabrication methods with enhanced control
over the porosity are the inverse opal hydrogelation which uses colloidal particles as templates for obtaining ordered macroporous scaffolds or cryogelation utilizing frozen solvent crystals as interconnecting
porogen [13].
3.2. Fibrous scaffolds
Fibrous scaffolds from different biodegradable polymers such as PCL
[14], PLA [15], PLGA [16], gelatin [17], cellulose [18], silk fibroin [19]
have been successfully prepared. Fibres possess the desired properties
as scaffolds for skin, cartilage, ligament, bone, muscle, vein and vehicles for drugs, DNA and proteins [20]. Nanofibres as scaffolds are
found to be more favourable biomaterials than micro fires because the
nanosize provokes cell to obtain typical in vivo morphology. Nanofibres
have the potential to provide guiding alignment for neurite growth by
establishing topographical clues affecting cell differentiation and fate
Fig. 1. A general scheme of various types of 3D scaffolds together with their
applications in tissue engineering.
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(from mm to nm) of geometrically complex objects, controlled pore size
and interconnectivity, and enhanced productivity of a wide range of
biomaterials. The computer-aided method allows creating scaffolds
incorporating patient-specific information in a uniquely designed
micro-environment. The data for tissue geometry can derive from
magnetic resonance imaging (MRI) or computed tomography (CT) that
allow for reconstruction in the 3D model (reverse modelling) [44].
Derived from the 3D medical image of the defect, this fabrication
technique is able to create patient-specific scaffold with the exact
geometry of the defect. The individualized scaffold design provides
production of more accurate and perfectly fitting implants together
with decreased fabrication time. However, the accuracy of the model
depends on the resolution of the device for image acquisition, 3D
printing technology, and machine used.
The additive manufacturing technologies could be classified into
scaffold-based (subdivided into cellular and acellular) and scaffold-free
3D printing [45]. Cellular 3D printing could be sub-categorized as extrusion-based, droplet-based, laser-based (that will be later explained),
and stereolithography (SLA) while acellular printing can be sub-divided
into SLS, selective laser melting (SLM), electron-beam melting (EBM),
fused deposition modelling (FDM), and melt electrospinning writing
[46]. SLA or vat polymerization is a process where a photoreactive resin
is selectively cured while a platform moves the scaffold after each new
layer is formed. The method is used for the synthesis of polymer,
ceramic and polymer-ceramic composite scaffolds with high accuracy
and resolution. Nonetheless, limited materials used as photoresist, residual toxic moieties, and need of post-processing remain challenges in
the biomedical application of SLA scaffolds. In contrast to SLS, SLM uses
a high energy density laser that completely melts the material thus
increasing the mechanical strength and surface quality of the scaffold.
The powders used require uniform distribution of spherical particles
with equal size. The smaller the size, the higher the resolution and
accuracy. When EBM is used, electron-beam gun in a high vacuum
selectively scans and melts only conductive metal powder (even with
high melting points) materials to form a 3D scaffold. The EBM-processed scaffolds have high surface roughness and limited accuracy [47].
FDM is an extrusion-based process that heats the material (polymer and
polymer-ceramic composites) before squeezing it out of a nozzle. By
moving the nozzle, the material is layer-by-layer deposited on a substrate. More comprehensive review regarding the methods of 3D scaffolds printing can be seen in Ref. [48]. Most printed scaffolds fulfil the
requirements for tissue engineering applications providing interconnected micro- and macroporosity, anisotropy, and heterogeneity.
Various types of materials could be used for printing of different scaffolds for skin, bone, nerve, muscle, cartilage regeneration or heart
tissue replacement [49,50]. Оnly CAD-based design could produce
scaffolds with regular and periodic structure with cube, diamond, gyroid, etc. unit cells. In that way, the porosity, pore size and surface-tovolume ratio can be exactly defined. However, biological materials
should be liquidized before printing and biologically relevant cellular
density is hard to be achieved [49]. Recently, the use of cryogenic 3D
printing shows great potential in the incorporation of high quantity of
bioactive moieties in situ into scaffolds. After fabrication at relatively
low temperature (−32 °C), high bioactivity and adequate mechanical
strength without post-treatment were achieved [51]. Nonetheless, the
cost of specialized equipment is still high for mass fabrication.
gradient microstructure formation [30]. However, the lack of specific
functional groups in synthetic nanofibres requires functionalization by
attaching different ligands (such as molecules, proteins, ceramics, etc.).
Mixed fibre mesh scaffolds provide the opportunity to develop constructs with enhanced properties (biomechanical, physicochemical and
biological) in comparison to the properties of individual scaffolds. For
example, when nanofibre-reinforced composites are used, higher mechanical strength than traditional unfilled composites could be obtained
[31]. Nonetheless, conventional fibre scaffolds suffer from some disadvantages like limited thickness and small pore size.
3.3. Microsphere/microparticle scaffolds
These scaffolds are intensively used in advanced applications such
as gene therapy, drug and growth factor delivery in a controlled fashion
[32] while demonstrating a certain degree of site-specific targeting
[33]. Bioactive moiety-delivery is determined by the ability of the
scaffold to control and trigger the substance release in a smart manner
through the interaction with bio-molecular stimuli [34]. Polymers with
low molecular weight demonstrate rapid drug release, whereas high
molecular weight microspheres achieve slower release [35]. Embedded
within 3D scaffold as building blocks, microspheres are capable of a
cumulative release of encapsulated bioactive substances [36]. Different
methods such as solvent vapour treatment, solvent/non-solvent sintering, oil-in-water dispersion, selective laser sintering (SLS) are used
for the production of microspheres and microsphere-based 3D scaffolds
[37]. The first three methods sinter polymeric microspheres at ambient
temperatures to make porous matrices for drug delivery or cell seeding.
The polymerization is conducted in heterogenous (two- or multi-phase)
systems where a monomer-rich phase is suspended in a solvent-rich
phase. During or after polymerization, aggregation, coalescence or
crosslinking of the formed microspheres/microparticles are used. The
SLS is an additive manufacturing technique for fabrication of complexshaped scaffolds without using moulds or preforms. The laser beam is
the heating source for selective sintering of the powder material (metal,
polymer, ceramic, composite) according to the predetermined model
geometry. An advantage of the sintered porous structure is the ability to
form a 3D porous structure suitable for bone regeneration [38]. For
example, lovastatin microparticles in polyurethane scaffold were found
to release lovastatin stimulating expression of BMP-2 growth factor in
osteoblast cells in 14 days period [39] thus combining better cellular
adhesion and controlled delivery of active biomolecules. However, cells
could not survive as long as they are more than a few hundred μm away
from blood vessels. Otherwise, gases, nutrients and soluble substances
should be supplied by slow diffusion which may cause delayed new
tissue formation or cell necrosis [40]. To encourage vascularization of
an appropriately designed scaffold, three main approaches are applied:
1) functionalization with VEGF, 2) seeding with endothelial cells, and
3) adding angiogenic gene carrying vectors or hypoxia mimic agents
[41]. Using the first approach, simultaneously loaded PLGA microspheres with BMP-2 and gelatin hydrogel with VEGF onto PP scaffold
have been shown to successfully induce both enhanced bone formation
and increased vascularization in rat bone in vivo [42].
3.4. Solid free-form scaffolds
In general, most conventional techniques for scaffold production are
incapable of creating complex scaffolds with precisely controlled microarchitecture and properties [43]. 3D scaffolds with controlled architecture and reproducible properties are created by stacking CAD/
CAM produced 2D shapes. After preparing 2D models, they are directly
transformed into STL files and transmitted to the additive manufacturing equipment. The whole 3D scaffold is obtained by successive
printing of 2D layers. The process includes precise X-Y-Z positioning
system, automated material-fed system, and computer-based software
for operational control. This allows high precision in spatial parameters
4. Types of 3D scaffolds based on the source material
Depending on the source of materials utilized for fabrication, 3D
scaffolds could be divided into alloplastic synthetic scaffolds and natural tissue scaffolds. Even though a wide range of materials evaluated
as hydrogels are commonly thought of in two groups, natural and
synthetic.
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chemistry and crosslinking) of hydrogels are crucial to ensure the
construct shape maintenance without compromising cell (cytocompatibility) or bioactive moiety spreading and function. Hydrogels are
currently used in cell scaffolds, bone regeneration, cartilage healing,
wound dress and drug or growth factor delivery [56]. The main challenge for the compatibility of hydrogels remains the toxic moieties or
chemicals used in the polymerization of synthetic or crosslinking of
natural hydrogel precursors, and other organic solvents, initiators,
stabilizers, etc. The methods used for fabrication of hydrogel scaffolds
include solvent casting/leaching, gas foaming/leaching, photolithography, electrospinning, 3D printing, etc.
4.1. Alloplastic synthetic scaffolds
Synthetically created 3D scaffolds are vastly used materials for
tissue engineering since they propose almost complete control over the
mechanical properties and architecture of the construct. Commonly
utilized materials are metals, bioceramics, glass and glass-ceramics,
synthetic polymers, and a myriad of composites. Overall, these synthetic biomaterials can be broken up into two major subgroups, nonbiodegradable and biodegradable. Recently, researchers emphasize the
use of biodegradable materials as they create a suitable microenvironment for growth of native tissue and are susceptible to cellular remodelling leaving in vivo produced natural matrix [52]. Based on their
application, the major groups of synthetic scaffold will be further discussed.
4.3. Natural tissue scaffolds
Natural tissues consist of cells, growth factors and extracellular
matrix (ECM). ECM is a heterogeneous hydrophilic 3D matrix that
provides a proper microenvironment for cell, accumulates and presents
growth factors, directs migratory cells, and participates in mechanical
signalling by mechanical receptors (integrins). The composition of ECM
usually includes vitronectin, fibronectin, collagen type I and II, decorin,
diglycan, laminin and perlecan [57]. The main components – collagen
fibres and proteoglycan (from protein and hyaluronic acid) filaments,
provide the durability and tensile strength of ECM. In order to exploit
the advantages of natural ECM, researchers are using the decellularization procedure to remove all cellular components that could cause an
inflammatory response in the host. Thus, the remaining ECM retains its
composition, architecture, integrity, biomechanical properties, biological activity, hemocompatibility and is able to direct the cell migration, tissue-specific gene expression, and to control cell fate. The process of removal of cells and their sources are crucial for the
decellularized ECM (dECM) performance. The decellularized material
could keep the whole organ intact or could be further processed by
cutting or digesting into the liquid to form a coat or ECM-containing
hydrogel. By using detergents or mechanical manipulations, natural 3D
scaffolds such as heart [58], lungs [59], urethra [60] and bladders [61]
are decellularized and functionalized by re-implanting host-specific
stem cells (re-cellularized) together with additional growth factors
[62]. The sources for creating ECM scaffolds include native devitalized
and decellularized human or animal (porcine, bovine) organs and tissues or de novo synthesized ECM from autologous, allogenic or xenogenic cells. Usually, human samples that could be used as dECM source
are either aged or diseased while xenogenic animal materials are potentially immunogenic. Moreover, difficulties in preparation could also
be faced. For in vitro synthesis (Fig. 2) of cell-derived dECM, a proper
selection of specific adult (induced pluripotent) or embryonic stem cells
should be made taking into account the medical application. It was
found that MSC-produced decellularized ECM dramatically increased
the growth and differentiation of neural cells [63] while maintaining
the multi-potentiality of MSCs during expansion in vitro [54]. Nonetheless, some disadvantages such as immunogenicity of material left or
inhomogeneity in cell distribution may appear, and risky and long
immunosuppressant treatment can be needed [64]. To avoid immunological response, autologous or allogenic sources are used. Using
autologous source cells, Hang et al. produced superior urea-extracted
dECM containing non-collagenous proteins enhancing MSC
4.2. Hydrogel scaffolds
Hydrogels are composed of hydrophilic polymer chains either
covalent or non-covalent (hold by intermolecular attractions) bonded.
When crosslinked through either covalent or noncovalent bonds, natural (gelatin, fibrin, alginate, agarose, etc.) or synthetic (PEG, PAA,
PEO, PVA, etc.) polymers form gels. In contrast to gels that are more
solid-like than liquid-like, hydrogels absorb large amounts of water and
swell without dissolving. In the swollen state, they are soft and elastic.
The inherent crosslinking (by using bi-functional monomers) of hydrogels allows them to retain 3D shape and swell without dissolving.
The higher the cross-linking extent, the lower the swelling. Both chemical and radiation crosslinkers have been used for the fabrication of
hydrogels [53]. Chemical hydrogels contain covalent bonds, whereas
physical hydrogels are maintained by ionic interaction, hydrogen
bonding or molecular entanglement between polymeric chains. Deriving from natural macromolecules, hydrogel scaffolds exhibit high
hydrophilicity, flexibility, biocompatibility, and degradability together
with limited mechanical properties, the difficulty of purification and
sometimes pathogen transmission and immunogenicity (depending on
the source). They can be also used as injectable materials able to adapt
the form of the damaged tissue. Synthetic polymers used for developing
hydrogels such as PEG, PLGA, PVA, PCL, PLA, PU, etc., propose tunable
and responsive physicochemical characteristics like modulus, water
affinity, degradation rate, etc. at the expanse of potential cytotoxicity
and lack of cell-adhesion moieties. According to their structure, both
natural and synthetic hydrogels could be amorphous or semi-crystalline, while depending on their response to environmental stimuli, hydrogels could be divided into conventional and smart (intelligent).
“Smart” hydrogels reversely change their swelling behaviour or structure in response to light, pressure, temperature, pH, ionic strength,
electric or magnetic field, and other stimuli that make them interesting
for the production of 4D scaffolds such as artificial muscles, self-regulating drug-delivery systems, etc. Hydrogels are also beneficial for cell
transplantation because they offer immuno-isolation while allowing
gaseous exchange as well as nutrients and metabolic substances to
diffuse [54]. They are important materials for scaffolds due to the
ability оf tailoring the mechanical properties, including peptide moieties preventing bacterial invasion, and adhering or suspending of cells
[55]. The appropriate rheological properties (dependent on hydrogel
Fig. 2. Schematic overview of in vitro preparation of ECM template.
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biomaterial characterized with good biocompatibility, tensile strength,
and corrosion resistance. Porous Ti scaffolds with 3D architecture
benefit the vascularization, nutrient and gas transport, and cell seeding
[70]. As previously explained, the pore sizes have a decisive influence
on bioactivity of the porous scaffolds. In order to determine the optimum pore size, Yang et al. fabricated screw-shaped Ti6Al4V dental
implant prototypes by laser beam melting (LBM) with three controlled
pore sizes (200, 350 and 500 μm) [71]. MC3T3-E1 cells showed improved attachment, proliferation and differentiation on both 350 and
500 μm pore size implants. Moreover, the 350 μm pore size alloy displayed the best mechanical stress distribution in the surrounding bone
under 20, 30, 40 and 80 N loading. When EBM processing was used,
Ti6Al4V scaffolds revealed better anti-corrosion ability with reduced
precipitates of harmful Al and V ions compared to wrought scaffolds
[72]. Moreover, the same EBM technology was capable of producing
double and triple-layered meshes from Ti6Al4V alloy [73]. The cylindrically shaped lattice scaffolds were composed of layers with different
porosity of 65–21% in order to mimic trabecular bone structure. The
graded (denser outer and less dense inner) tubular design of the scaffold
demonstrated decreased Young moduli (from 0.9 to 3.6 GPa) compared
with a dense alloy that brought it closer to that of human bones. By
changing the morphology of the lattice structure, the deformation behaviour of the scaffold was able to provide a unique combination of
high ductility, energy absorption, and strength simultaneously [73].
Nonetheless, comparing the performance of EBM-produced Ti6Al4V
and Co–Cr scaffolds with similar porosity of 67–70% and pore size of
470–545 μm on bone tissue growth, Shah et al. [74] established higher
osteocyte density at the periphery of Co–Cr scaffold. The authors explained that fact with the existence of more favourable biomechanical
environment and higher stiffness within Co–Cr alloy as opposed to
Ti6Al4V. Fousová and her group produced fully interconnected open
porous scaffold with rough surface by adhering spherical particles
(15–50 μm) from 316L stainless steel using SLM [75]. The mechanical
properties in compression of the scaffold with square-shaped pores with
size of 750 μm and porosity of 87 vol% approached those of human
trabecular bone. In another study, pre-cultured cell-loaded with rat
bone marrow stromal cells (BMSCs) Ti fibre mesh with volumetric
porosity of 86% and fibre diameter of 45 μm demonstrated an almost
complete absence of inflammatory cells and improved bone healing
capacity [76]. Although the bone defect symptoms were improved, the
use of embedded BMSCs requires appropriate donor sites that could
cause additional pain and infection. The main problems with titanium,
Co–Cr and stainless steel scaffolds remain the lack of metabolization
over time, need of repeated surgery, risks of ion release and abrasion
(that could trigger inflammatory cascade), disease infection, low tissue
adherence and prolonged recovery time.
Porous Mg scaffolds are promising biomaterials for bone substitute
application owing to their good mechanical properties and biodegradability [77]. Moreover, released Mg ions and biodegradation products
in endosseous sites demonstrated the ability to induce new bone formation and neovascularization [78]. In this context, open porous Mg
scaffolds with 250 and 400 μm pore size and 55% porosity were found
to exhibit good cytocompatibility and enhanced alkaline phosphatase
(ALP) activity indicating osteogenic differentiation in vitro whereas in
vivo the scaffold with larger pore size promoted vascularization and
higher bone mass formation in a rabbit model [79]. The compressive
strengths were about 41 and 46 MPa whereas Young's moduli reached
values of 2.2 and 2.4 GPa for 250 and 400 μm pore size scaffold, respectively which values exceeded Young's moduli of the cancellous
bone. However, weak points of porous Mg scaffolds remain the fast
degradation rate and accelerated micro galvanic corrosion by the presence of impurities [80,81] together with the production of a large
amount of hydrogen gas during in vivo degradation [82]. In addition, it
was found that Mg ions can negatively affect the function of red blood
cells and hemolysis ratio in direct contact with blood [83].
proliferation, migration and differentiation [65]. This strongly suggests
that the origin and preparation of ECM predetermine the biological
activities of biomaterial scaffolds. Additionally, different cell types
could be mixed to create gradient tissue scaffolds. The cells plated in 3D
scaffolds secrete and integrate new components to form ECM trying to
shape their environment. Nevertheless, these artificially fabricated
dECM scaffolds are less dense and have restricted chemical complexity
because their organization is a result of spontaneous, rather than celldirected polymerization [66].
Once implanted, the dECM releases soluble peptides while degrading that are able to chemo-attract stem cells to the injured sites
[67]. These chemoattractive properties and degradation products vary
depending on the age, species, gender or physical characteristics of
ECM source. For example, a cell-derived ECM was obtained from cells
grown in 3D cultures of fetal and adult synovium-derived stem cells
(SDSCs) that were afterwards decellularized for ECM scaffold fabrication [68]. The fetal ECM provided a better microenvironment for proliferation, while adult ECM was advantageous for chondrogenic and
adipogenic differentiation. However, all these costly and time-consuming procedures restrict the clinical application of the cell-derived
matrix scaffolds.
It follows that an important factor that determines the biocompatibility of the scaffold except for the chemistry, morphology, structure,
and properties, is the processing of a biomaterial. To overcome the
disadvantages of a certain fabrication method, a combination of other
methods could be used that complicates the existing production processes. Considering the variety of biomaterials used, not all of them are
suitable for a given fabrication technology. This is a reason for constant
modification of the materials which broaden their use. Since the surface
characteristics such as wettability, chemistry, charge, and surface
roughness are also able to regulate the contacts with the living cells and
ECM proteins, a variety of surface treatments are proposed to optimize
the biocompatibility of the scaffold while sealing undesirable additives
and regulating absorption or corrosion rate. The next paragraphs discuss the properties, advantages and disadvantages of the major types of
biomaterials used in the fabrication of 3D scaffolds dedicated for various applications.
5. Scaffolds for hard tissue application
Bone tissue consists of organic (predominantly collagen matrix,
22 wt%), inorganic (mainly HAp crystals, 69 wt%) components and
water (9 wt%) [69]. In its hierarchical organization two main types of
bone structures could be distinguished: trabecular (porous) and cortical
(compact) bone, both reinforced with collagen fibres. Because of the
natural self-repairing and regeneration ability of bones, most of the
produced scaffolds aim at providing regenerative signals to osteogenic
cells to enhance regeneration and repair [69]. A major difficulty in
designing scaffolds for load-bearing application is to simultaneously
tailor all biomaterial requirements that are competing in nature. Hard
tissue scaffolds should not only be biocompatible, well-integrated with
the native tissue, and easily produced but should also have an ideal
replacement rate and highly porous structure which doesn't significantly compromise the adequate mechanical properties. Commonly
used alloplastic synthetic scaffolds as bone grafting materials that
mimic the native bone tissue and provide structural and mechanical
support are metals, biomimetic ceramics and composites.
5.1. Non-biodegradable and resorbable metal scaffolds
As first-generation materials for bone substitutes, metallic 3D scaffolds are largely popular for load-bearing applications compared with
ceramics or polymers because of their high mechanical strength, fatigue
resistance, and printing processability. Commonly used metallic biomaterials include titanium, stainless steels, cobalt-chromium (Co–Cr)
based alloys, and magnesium (Mg). Titanium (Ti) is a metallic
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47 MPa compressive strength value or 30% improvement over pure
SiO2 scaffold, no cytotoxicity, and good biocompatibility when tested in
MG-63 osteoblast-like cells in vitro. Pressure extruded ceramic composite scaffolds of calcium sulphate hydrate (CSH) and mesoporous
bioactive glass (MBG) with different mass ratio of the components,
uniform square macroporous structure (67–68% porosity) and pore size
of 350 μm indicated stimulated adhesion, proliferation and osteogenicrelated gene expression of hBMSCs (human bone marrow-derived stem
cells) in vitro [91]. In vivo results in rats demonstrated that the scaffolds
promoted new bone formation in calvarial defects compared to pure
calcium sulphate hydrate (CSH) scaffolds.
Application of growth factors or cytokines is used to promote cell
differentiation, tissue formation or neoangiogenesis. On one hand, most
of the growth factors (GFs) have a short half-life in circulation and on
the other, fast and uncontrolled release of GFs may increase the adverse
effects on non-target sites. For that reason, the controlled release of
these substances is a desirable capability of the scaffold. To facilitate
vascularization, Li et al. printed at room temperature (without sintering
afterward) mesoporous with approximately 300 μm macropores silica/
calcium phosphate cement (CPC)-based 3D scaffold with concomitant
Si ion (promoting vascular tissue ingrowth) and recombinant human
bone morphogenic protein-2 (rhBMP-2) release stimulating osteogenesis of human bone marrow stromal cells [92]. The produced scaffolds
with uniform interconnected pore structure induced osteogenic differentiation of hBMSCs and vascularization of human umbilical vein endothelial cells (hUVECs) in vitro. Implanted in femur defect rabbit
model abundant new vessels around the complex scaffold and rapid rate
of osteogenesis were observed as compared to CPC control scaffolds
[92]. Moreover, various studies demonstrated improvement in biological and physicochemical properties of mesoporous bioactive glass
(MBG) scaffolds by incorporation inorganic therapeutic components
such as Sr [93], Zn, Mg [94,95], Zr [96], Ga [97], or B [98]. Incorporation of beneficial Fe (5–10%) in MBG turned the scaffolds with
high porosity (83%) into biocompatible magnetic material that exposed
to the external magnetic field had great potential for use in hyperthermia therapy (at around 43 °C) of malignant bone tumours [99].
The scaffolds with a hierarchical structure of large 300–500 μm pores
and 4.5 nm mesopores demonstrated a slight decrease in compressive
strength (46–48 kPa) as opposed to pure PBG scaffolds (50 kPa). The
multifunctionality of Fe-MBG scaffolds that makes them suitable for
therapy of malignant bone diseases is schematically presented in Fig. 3.
Five percentage of Fe in MBG scaffold benefited the mitochondrial activity and gene expression of BMSCs indicating improved osteoconductivity of the implant. This fact the authors explained with
changes in ionic composition and pH value that improved cell viability
and differentiation [99]. Moreover, mesoporous silica displayed additional capability of local drug release owing to its large nanopore volume, size and surface area. To demonstrate that Zhang et al. fabricated
5.2. Low degradable bioceramic, glass and glass-ceramic scaffolds
Ceramic biomaterials usually include inorganic calcium or phosphate salts that have osteoconductive (promote new bone ingrowth)
and osteoinductive (promote osteoblastic differentiation) properties.
Ceramics could be classified into inert (non-absorbable), semi-inert
(bioactive) and non-inert (resorbable) [84]. They are commonly brittle
in nature but show good compression and corrosion resistance. Hydroxyapatite (HAp, Ca10(PO4)6(OH)2), β-tricalcium phosphate (β-TCP,
Ca3(PO4)2) and bioactive glasses are among the most common biomaterials used for 3D scaffolds in bone regeneration. Except conventional
methods for production of porous ceramic scaffolds such as polymer
sponge method, salt leaching, dual-phase leaching, gel casting, etc., the
major techniques available for ceramics 3D printing are: 1) agglomeration with polymer, chemical/physical solidification and thermal
processing, and 2) sacrificial inverse matrix printing, infiltration with
ceramic slurry and burning the negative [85].
Because of their similarity to bone, bio-resorbability and good biocompatibility, calcium phosphate biomaterials are widely used for orthopaedic and dental implant applications. 3D printed β-TCP scaffolds
showed an increase in uniaxial strength with the increase of sintering
temperature and time [86]. The compression strength could be additionally improved by controlling particle size distribution and binder's concentration. Another approach proposed by Deng at al. includes
incorporation of Mn in β-TCP scaffolds that improved both biomechanical (density and compressive strength) and biological properties [87].
The ionic products from the scaffold promoted the proliferation of
rabbit chondrocytes and rBMSCs in vitro and improved regeneration of
subchondral bone and cartilage tissue as compared to β-TCP scaffolds
upon transplantation in rabbit models. However, similarly to HAp, the
application of β-TCP as bone regeneration scaffold is limited because of
its overall poor mechanical properties. For that reason, 3D printed
porous β-TCP scaffolds with interconnected squire channels (500 μm in
size) were mechanically strengthened by SiO2 (silica; 0.5 wt%) and ZnO
(0.25 wt%) dopants. The cylindrical scaffolds demonstrated increased
density and a mean increase in compressive strength of 250% after
sintering at 1250 °C when compared to pure β-TCP scaffolds [88]. A
greater rate of attachment and proliferation of human fetal osteoblast
cells (hFOBCs) within the doped scaffolds was also observed.
In general, bioactive glasses (BG, CaO–SiO2–P2O5 composition)
share better degradation properties and bioactivity than HAp and βTCP. Therefore, the biodegradation performance of ceramic bone implants could be improved by introducing mesoporous silica-based particles that also promote Si ion release known to be essential for proangiogenesis and healthy bone development [89]. Bioceramic scaffolds
with 34% porosity were 3D printed with laser-aided gelling from
CaCO3/SiO2 (5:95 wt%) sol in the form of cylindrical specimens and
sintered at 1300 °C afterwards [90]. The bioceramic scaffold showed
Fig. 3. A scheme illustrating the potential application of Fe-MBG scaffolds for malignant bone treatment (hyperthermia) and regeneration of the defect bone.
Adapted from Ref. [99].
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M.P. Nikolova and M.S. Chavali
3D printed porous Sr-containing MBG porous (400 μm pore size, 70%
porosity) scaffold that showed sustained dexamethasone (anti-inflammatory drug) release with a rate depending on Sr dissolution
characteristics of the scaffold [100]. To combine drug-delivery properties with additional angiogenetic and antibacterial properties, multifunctional Cu-containing MBG scaffold with large pore size of
300–500 μm loaded with ibuprofen was prepared by simple polymer
sponge method [101]. The ion (Cu2+) release was able to induce a
hypoxic cascade of hBMSCs activating neovascularization, promote
osteogenic differentiation, and prevent osteomyelitis incidence by infections. In this respect, composite ceramic scaffolds could meet all the
requirements of a bioresorbable therapeutic cell and drug (or antibiotics and growth-factors) carriers for improved healing in bone tissue
engineering. The major disadvantages of ceramic scaffolds are their
brittleness, compressive strength lower than that of the human bone
(100–230 MPa [102]) at high volume percentage of porosity, and the
need of high sintering temperature that limits the incorporation of
bioactive molecules in the scaffolds. Some ceramics such as calcium
sulphate hydrates (CSH) display faster degradation rate than the formation of new bone and a release of acidic degradation products that
are not beneficial for cell proliferation and viability [103]. Other
ceramics like wollastonite (CaSiO3) are difficult to be cut and shaped in
uniform porous scaffold with controllable mechanical properties and
porosity [104]. Moreover, the highly promising MBG scaffolds could
have high degradation rate together with unstable interface/surface
performance that could also decrease cell attachment and proliferation.
encapsulation of TGF-β1 with a higher dose of BMP-2, the histological
analysis demonstrated tissue repair after very short (2 weeks) period,
higher quality cartilage with improved surface regularity, and very
good tissue integration in the rabbit model. The controlled release rate
of GFs from the scaffolds preserved cartilage integrity from 12 to 24
weeks [112]. Using non-biodegradable orientated in parallel arrays of
polystyrene sub-micron fibres with diameter (630–760 nm) approximating the diameter of ECM fibres, coated with fibrin and bioprinted
with tendon- or bone-promoting-GFs (fibroblast growth factor-2 (FGF2) or BMP-2, respectively), Ker et al. [113] inhibited the myocyte differentiation of mouse C2C12 myoblasts and stimulated cell alignment,
tenocyte and osteoblast differentiation. The bioprinting technique gives
the possibility to have direct control over the concentration of the immobilized growth factors. Therefore, using hydrogel with embedded
GFs, it is possible to present a precise quantity of extrinsic factors and
direct cells rather than adding them in the culture medium.
5.4. Particle loaded polymeric and other composite scaffolds
The development of composite scaffolds allows for the engineering
of biomaterials with suitable mechanical and physiological properties
by controlling the type, size, fraction, morphology and arrangement of
the reinforcing phase. Moreover, the degradation behaviour of the
scaffolds could be altered by adding bioactive phase in the matrix
[114]. For instance, PCL is relatively elastic and drug permeable but has
poor mechanical stability (low modulus with high elongation behaviour), slow degradation (2–4 years) in living tissue, and poor cellular
adhesion properties. However, it shows blend compatibility with other
biomaterials such as Sr-hydroxyapatite [115], nanohydroxyapatite,
nanocellulose, carbon nanotubes [116,117], cyclodextrins [118], etc. It
follows that PCL possesses the capacity for functionalization while
overcoming the poor mechanical stability and low bioactivity of the
unmodified polymer. PGA/HAp composites demonstrated high bioactivity, osteoconductivity uniform cell seeding, cell ingrowth and tissue
formation as hard tissue scaffolds [119]. By releasing Ca and Si ions,
MBG powders were also found to neutralize the acidic degradation of
synthetic polymers such as PLGA in composite scaffolds and stimulate
thereby cell response [120]. The ceramic mesoporous powders increased hydrophilicity, water absorption and degradation rate of the
composite compared to pristine PLGA scaffolds. Electrospun hybrid
scaffolds of randomly orientated fibres (average diameter 4.11 μm) of
poly (3-hydroxybutyrate-co-3-hydroxyvalerate) and silicate containing
hydroxyapatite nanoparticles (PHBV-SiHAp) seeded with hMSCs
showed the largest adhesion and differentiation ability compared with
pure piezoelectric PHBV and non-piezoelectric PCL scaffolds [121]. The
composite scaffold revealed improved mineralization and superior osteoinductive properties either because of changed scaffold chemistry by
the presence of bioactive SiHAp nanoparticles or by altering scaffold
charge due to inherited piezoelectricity of PHBV.
In another study, calcium phosphate cement (CPC)-based scaffold
incorporating chitosan, absorbable fibres, and hydrogel microbeads
possessed 4-fold increased flexural strength and 20-fold enhanced
toughness compared to rigid CPC microbead scaffolds which matched
the strength of cancellous bone [122]. These non-rigid scaffolds demonstrated excellent proliferation, osteodifferentiation and enhanced
mineralization when incubated with human umbilical cord mesenchymal stem cells (hUCMSC) in vitro. Similarly, freeze-dried porous
fluorhydroxyaptite (FHAp)-Mg-gelatin scaffold with different amount
of FHAp (30, 40, 50 wt%) and pore size of 150–250 μm improved the
compressive strength of the polymer from about 2.1 MPa to 3.3 MPa at
the highest amount of FHAp and supported the proliferation and adhesion of MG-63 cells for bone regeneration [123]. Graphene alone was
found to stimulate osteogenesis [124] while covalent associations of
graphene oxide (GO) flakes (4 nm thick) in porous collagen scaffolds
directed the stem cells fate toward osteogenic differentiation by upregulating cell-adhesion molecules and stiffening the scaffold [125]. As
5.3. Polymer scaffolds
Polymer materials are vastly used to renovate the traumatized tissue
due to their unique properties such as biocompatibility, reproducible
mechanical, physical properties, workability, and low price. Because
non-biodegradable synthetic polymer scaffolds require a procedure of
surgical removal, they are not widely used. Some acrylic polymers like
PHEMA, PHPMA, PMMA [105,106], and conductive polymers [107]
are such representatives. Two-part self-polymerizing PMMA cement is
known to be the most enduring material in orthopaedic surgery used in
total joint replacement for fixation of components [108]. However,
drawbacks such as aseptic loosening caused by monomer-mediated
bone damage, inherit inert properties, and mechanical mismatch during
long term wearing are also reported [108].
Biodegradable polymers with tunable degradation rates could be
both natural (from human or animal tissues) and synthetic. Their degradation rate depends on molecular weight, the structural arrangement of macromolecules (amorphous/crystal structure), isomeric
characteristics, formulation, architecture, and quantity of the material
[109–151]. Natural polymer scaffolds usually demonstrate a lack of
immune response and better cell interactions while synthetic polymers
are cheaper, stronger, and have better functionality although sometimes triggering immune response and toxicity [37]. Biodegradable
polymers demonstrate suboptimal load-bearing capacity when used
alone that limit their application as hard tissue scaffolds. However, in
pulpodentinal complex and periodontal apparatus, the smaller sizes and
difficulties in reaching the target places require soft, injectable scaffolds
that match irregular patient defects [110]. Moreover, the ability of
polymers to incorporate diverse bioactive moieties could be most
eminent to induce osteogenic differentiation in less loaded constructs.
For example, immobilizing or including GFs within the scaffold are
promising approaches as many GFs exhibit inherited binding properties
to molecules of ECM through specific binding protein intermediaries
[111]. Bilayer system of porous PLGA cylinder overlaid with PLGA
microspheres dispersed in alginate sponge matrix was examined for
localized delivery of pre-encapsulated transforming growth factor-β1
(TGF β1 – 50 ng) and bone morphogenic protein-2 (BMP-2 – 2.5 and
5 μg) applied for osteochondral defect repair [112]. In vitro, the total
dose of GFs was delivered at the end of the sixth week. After co279
M.P. Nikolova and M.S. Chavali
Table 2
Composite scaffolds for bone tissue application.
280
Biomaterial composition
Fabrication
Cell type
Outcome
Ref.
Gelatine, alginate, HAp scaffolds
Extrusion
hMSCs
[134]
Chitin-nanoHAp scaffolds
Gelatin-carboxymethyl chitosan-nanoHAp scaffolds
Glycol chitosan-hyaluronic acid-nanoHAp scaffolds
Chitosan, gelatin, and GO containing scaffolds
Freezing/thawing method
High stirring-induced foaming
and freeze-drying
Injectable
Freeze-drying
Chitosan-nanoHAp containing Cu/Zn alloy nanoparticle scaffolds
Freeze-drying
COS-7 (fibroblast-like) cell line
Human Wharton's jelly-derived
mesenchymal stem cell microtissues
MC-3T3-E1
Rat calvarial osteoprogenitor cells and
mouse mesenchymal stem cells (C3H10T1/
2)
Rat osteoprogenitor cells
Cell survived the printing process and showed 85% viability
after 3 days
Good adhesion and proliferation of cells
Cell growth, proliferation and differentiation; high
mineralization capacity
Cytocompatibility with cells well attached to the pores
Promote differentiation into osteoblasts; increased collagen
deposition in vivo
[139]
Blended PLGA-silk fibroin fibrous scaffold coated with HAp
Electrospinning
MSCs
Micro-nano PLGA-collagen – nanoHAp rods scaffolds
Alginate-PVA-HAp hydrogel scaffold
Electrospinning
Bioprinting
MC3T3-E1
MC3T3
Tri-layer scaffold consisting of superficial PVA/PVAc-simvastatin (a type of
statin)-loaded layer, followed by PLC-cellulose acetate-β-TCP layer and final
PCL layer
Laminated nanoHAp layer on PHB (polyhydroxybutyrate) fibrous scaffold
Electrospinning
MC3T3-E1
Increase protein adsorption and antibacterial activity; no
toxicity towards osteoprogenitor cells
Increased adhesion, proliferation and differentiation towards
osteoblasts; excellent cytocompatibility and good osteogenic
activity
Improved osteogenic properties; bioactivity
Excellent osteoconductivity; well distributed and encapsulated
cells
Higher mineralization; enhanced cell attachment and
proliferation
Electrospinning
MSCs
[144]
PMMA-nHAp decorated cubic scaffold
Solvent casting and particle
leaching
Emulsion and solvent evaporation
method and sintering
Electrospinning
MG-63
G-292 cell lines
Better adherence, proliferation and osteogenic phenotype
formation
Friendly environment for cell growth and protection from
microbial infection
Increased cell viability; a higher amount of bone formation
[146]
MC3T3-E1
Scaffolds promote osteogenic differentiation
[147]
Electrospinning
Osteocytes
[148]
Self-assembly
Electrospinning
rBMSCs (rat bone mesenchymal stem cells)
mBMSCs (mouse bone marrow stem cells)
Retard corrosion and increased osteocompatibility; higher cell
attachment and proliferation
Enhanced proliferation and osteogenic gene expression
Faster proliferation; early stage osteogenic differentiation
PLGA/TiO2 nanotube sintered microsphere scaffolds
PU fibrous scaffolds loaded with MWCNTs (0.4 wt%) and ZnO nanoparticles
(0.2 wt%)
PCL-nanoHAp nanofibre layer deposited on Mg alloy scaffold
Porous rGO-nanoHAp scaffold
PLLA - osteogenic dECM (from MC3T3-E1) scaffolds
[135]
[136]
[137]
[138]
[140]
[141]
[142]
[143]
[145]
[149]
[150]
Bioactive Materials 4 (2019) 271–292
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M.P. Nikolova and M.S. Chavali
opposed to pure collagen scaffolds (14.6 kPa), GO-collagen construct
showed elastic moduli of about 39 kPa. It was also demonstrated that
adding silk (2.5 and 5 wt%) to mesoporous bioactive glass (MBG)
scaffold (94% porosity) improved its mechanical properties, cell attachment, proliferation and controlled drug release [126]. The compressive strength rose up from 60 kPa up to 250 kPa when incorporating
5 wt% silk in MBG scaffold which means 300% increase. The degradation rate of the composite was slower as opposed to that of a MBG
scaffold and a thin silk film on the pore walls (200–400 μm pore size)
was formed. Thus formed stable surfaces with fibroin provided better
support for BMSCs proliferation and differentiation [126].
In a remarkable study, Cheng et al. tried to construct highly porous
bilayer scaffold with interconnected honeycomb structure with chondral (consisting of plasmid TGF-β1-activated chitosan-gelatin scaffold)
phase integrated into osseous (of plasmid BMP-2-activated HAp/chitosan-gelatin scaffold) phase [127]. This bilayer gene-activated scaffold
with pores of about 50–100 μm was dedicated to spatial control of localized gene delivery that could induce the mesenchymal stem cells in
different layers to differentiate into chondrocytes and osteoblasts in
vitro and in vivo. The spatially controlled and sustained delivery of
plasmid DNA encoding for tissue inductive factors maintained the
specific cell phenotypes of complex osteochondral integrated tissue
derived from a single stem/progenitor cell source in vitro. In vivo, the
complex scaffold gave simultaneous support of articular cartilage and
subchondral bone [127].
The surface properties of 3D scaffold biomaterials could be altered
in order to promote their bioactivity, corrosion resistance, hydrophilicity, mechanical strength or tribological properties. For example,
bioactive adhesive molecules such as collagen, fibronectin, growth
factors, insulin, etc. can be covalently or physically attached on the
biomaterial surface. These modified 3D scaffolds are able to modulate
in a complex fashion the cellular response. Such hybrid polymericbioceramic porous (over 60% porosity) scaffolds of β-TCP/HAp vacuum
coated with alginate demonstrated improved osteoblast adhesion, maturation and proliferation as well as an enhanced mechanical performance involving increased fracture toughness, Young's modulus, and
compressive stress close to that of cancellous bone [128]. The improved
osteoblast adhesion and migration the authors attributed to the additional biocompatibility created by alginate via changing surface microtopography and roughness. Another approach aiming at increasing the
bioactivity and biocompatibility of metal implants is deposition of
wollastonite (CaSiO3) [129], calcium phosphate [130], biomimetic
nanoapatite [131], or polycrystalline diamond [132] coatings on different titanium scaffolds. When EBM-manufactured Ti6Al4V scaffold
was functionalized with Ag and CaP nanoparticles via electrophoretic
deposition, the change in hydrophobicity and surface roughness at
nanoscale provided physical cues that disrupted bacterial adhesion
[133]. The silver nanoparticles at concentration 0.02 mg/cm2 in CaP
film covered with positively charged PEI indicated bacteriostatic activity against gram positive bacteria, S. aureus, during 17 h of exposition. Although bi- and multiphasic scaffolds provide enhanced
mechanical and biological performance, there is always a concern
about the adhesive strength between the adjacent layers or contacting
non-homogeneous biomaterials that could lead to delamination or
crack propagation. Although improving the surface performance, these
additional modifications are usually cost- and time-consuming.
A summary of recent studies on composite scaffolds using numerous
production technologies and stem cells dedicated to a hard tissue application is reported in Table 2.
6. Scaffolds for soft tissue application
By careful selection of biomaterials and key scaffold characteristics,
researchers developed novel techniques for developing complex architectures with desired properties for soft-tissue engineering applications.
These scaffolds should be fabricated so that to regenerate and mimic
both the anatomical structure and function of the original soft tissue to
be repaired. Polymers are commonly used biomaterials to construct soft
matrices that are widely utilized for the production of most of the
transplanted organs such as kidneys and liver [2], but also successfully
applied for muscles, tendon [151,152], heart valves, arteries [153],
bladder and pancreas [154,155] regeneration.
6.1. Synthetic polymer scaffolds
Synthetic polymers are easily produced under controlled conditions.
Their physical and mechanical properties are tunable, predictable and
reproducible and thus, could be tailored for the production of scaffolds
dedicated for a specific application. The most commonly utilized synthetic polymer for drug, gene, and growth factor delivery application
[156–158], is PLGA, a flexible and permeable copolymer of lactic and
glycolic acid. Each lactic acid residue includes a pendant methyl group
which makes the surface hydrophobic. The polymer chains show a lack
of functional groups and biodegrade in non-toxic but highly acidic
glycolic and lactic acid [159]. To address the problem with acidity, a
greater amount of glycolic than lactic acid could be used to form PLGA
that lessens the degradation rate and lowers the acidic byproducts.
Because of their semi-permeability, PLGA microspheres were used to
entrap living cells like myoblasts [160]. When subcutaneously implanted, the former simultaneously released therapeutic anti-inflammatory drug to protect these cell from the host immune system
(Fig. 4). The myoblast cells retained their viability in 30 day period
suggesting that the semi-permeable microspheres allowed appropriate
diffusion of nutrients and oxygen. These encapsulated cell implants
loaded with dexamethasone, after 60 days of subcutaneous post-implantation in rats diminished the inflammatory response by sustain
delivery of the drug. The histological analysis revealed that blood capillaries surrounded the microcapsule aggregates in 100 μl dexamethasone-treated mice group [160].
Not only encapsulated but also aligned nanofibre scaffolds were
found to be suitable for highly organized soft tissues such as skeletal
muscle, ligament and peripheral nerve regeneration whereas random
Fig. 4. Schematic representation of subcutaneous microenvironment after implantation of encapsulated myoblast cells and microspheres releasing dexamethasone in
mice. Adapted from Ref. [160].
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nanofibres were useful for cartilage and skin regeneration application
[161]. Since muscle fibres must grow parallel to one another following
identical anisotropy, electrospun PLGA scaffolds of aligned fibres with a
diameter of about 0.6–0.9 μm were found to provide topographical cues
guiding the alignment of myoblast cells (C2C12 murine myoblasts) and
encouraging their differentiation as compared with randomly oriented
fibre substrates [162]. Moreover, the incorporation of polyaniline in
PCL well-ordered fibrous scaffolds increased the electrical conductivity
of constructs thus giving not only topographical but also electrical cues
to C2C12 myoblast cells both synergistically improving myotube maturity [163]. Another approach applied is using permeable core-shell
nanofibres that incorporate bioactive agents and enable controlled release within the tissue. The active substance is usually entrapped in the
core layer. Li et al. used co-axial electrospinning where two syringe
pumps fed both solutions separately (Fig. 5) to produce poly(L-lactideco-ε-caprolactone), (PLC) fibrous scaffold of different LA:CL ratio and
BSA (model protein) containing core, namely PLC(50:50)BSA and
PLC(75:25)BSA [164]. The BSA release of PLC(50:50)BSA and proliferation rate of hMSCs toward smooth muscle cells were higher than
those of PLA(75:25)BSA. Except for tunable mechanical properties,
these core-shell nanofibres possessed regulable release potential as
medical 3D scaffolds.
Damaged neural tissues are known to regenerate for a longer period
of time whereas sometimes larger nerves never recover. Similarly to
muscle scaffolds, for enhanced nerve regeneration fibrous scaffolds of
PLLA, PCL, PGA, etc. with aligned or random arrangements are commonly used [165]. Moreover, different studies demonstrated that
therapeutic cell-laden scaffolds also enhanced in vivo soft tissue regeneration. For example, by combining longitudinal aligned electrospun scaffolds of PCL and PLGA (proportion of 4.5:5.5) with hDFCs
(human dental follicle cells), neural regeneration was stimulated [166].
The authors found no cytotoxic reactions and when transplanted to
restore the defect in rat spinal cord, re-myelination and induced tissue
polarity occurred. It was also reported that incorporating of nerve
growth factor (NGF) into aligned core-shell PLGA nanofibres enhanced
physical and biomolecular signalling and promoted better nerve regeneration in 13-mm rat sciatic nerve defect than only PLGA scaffold
[167]. Similarly, a strong electrospun fibre-aligned scaffold of PCL and
PLLA with incorporated bFGF and PGF (platelet-derived growth factor)
was found to up-regulate the gene expression of hMSCs in vitro [168]
critical for the viability and repair of the anterior cruciate ligament that
had a poor healing capability.
The complexity of mechanical requirements and physical properties
of cartilage hinders the fabrication of effective artificial scaffolds for
cartilage regeneration. Because of low coefficient of friction (0.02–0.05
against smooth and wet substances), high permeability to fluids, and
biocompatibility, PVA-based hydrogels were developed to be used as
synthetic scaffolds for articulating cartilage [169]. Kim and his group
developed macroporous PVA sponge incorporating in its pores encapsulated rabbit chondrocytes in photocrosslinkable PEG derivates
that improved in vitro chondrocyte functions and collagen accumulation
within the scaffold [170]. In an ectopic mouse model, the mechanical
properties of the cell-laden PVA-based scaffold were biomimeticaly
reinforced by over 80% compared to their acellular counterparts. Similarly, by encapsulating chondrocytes in a photopolymerized degradable PEG tiol-ene hydrogel scaffold with localized presentation of
TGF-β1, Sridhar et al. demonstrated in vitro cell viability, proliferation,
and cartilage-specific molecules generation at a higher rate than
without GF [171]. It follows that tethering GFs into synthetic polymer
scaffolds integrates promoting effect on cells and provides advantages
for clinical applications. An issue when using extrusion printing technology for scaffold production is the elevated temperature (above
100 °C) that prevents the incorporation of bioactive materials promoting the healing process. Guo et al. proposed lowering the printing
temperature of PLGA scaffold by utilizing dimethyl sulfoxide (DMSO) as
a solvent that allows for the incorporation of proteins favourable for
induction of hMSC differentiation [172]. After the solvent treatment,
the material became tougher with improved flexibility and compressive
strength similar to that of native cartilage while the activity of the
growth factors was retained with the cold printing method used.
In another study, thermoplastic polyurethane (TPU) (25%)/PCL
(75%) blends with thermally induced shape memory were produced by
melt blending and showed 98% shape fixing ratio and 90% shape recovery ratio upon heating (60 °C in hot water bath). 3T3 fibroblasts
cells cultured on the TPU/PCL scaffold indicated high viability together
with obvious cell-substrate interactions [173] that make the composite
biomaterial suitable for surgical sutures or other medical devices.
However, many synthetic polymers demonstrate low cell-affinity surfaces [174], hard to control degradation rate, and in vitro cytotoxicity
[159]. Other problems encountered are lack of hydrophilicity, bioactivity, and release of acidic by-products that lower pH values and
trigger inflammatory response [175].
Synthetic supramolecular scaffolds are also applied in regenerative
medicine. Synthetic peptides usually form a network of waving nanofibres with a diameter of around 10 nm. They successfully mimic the
native microenvironment of ECM because of their biomechanical
properties and nanoscale network. Self-assembly synthetic peptidebased hydrogels are attractive candidates for scaffold materials showing
biocompatibility and ability to attach versatile chemical groups. They
also demonstrate hypo-immunogenicity, slow degradation (by
Fig. 5. A scheme illustrating the principle of co-axial electrospinning where the polymer in a solvent coats the inner aqueous solution while immerging from the
needle. As a result, a smooth and beadless core-shell nanofibre is formed. Adapted from Ref. [164].
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adhesion of natural killer cells [190] that are responsible for activation
of stem cell migration.
The good gelling properties of alginate (polysaccharide extracted
from brown algae) allow the polymer to be used as an injectable scaffold for damaged cartilage repair. When human dental pulp stem cells
cultured in 3% alginate hydrogel were implanted in a rabbit cartilage
defect, significant cartilage regeneration was observed after 3 months
[191]. Suitable not only for cartilage [192] and tendon [193] regeneration but also for wound dressing application [194] is silk fibroin,
a protein distinguished by its high tensile strength and enhanced cellular adhesion properties. Hyaluronic acid (HA) is a linear polysaccharide that is one of the most important constituents of ECM. Due
to the encapsulating and swelling capability of HA, it is used for delivery application [195] whereas porous HA scaffolds assembled by
layer-by-layer spray assisted method demonstrated enhanced in vitro
human keratinocytes adhesion and proliferation [196]. However, native HA does not support cell adhesion and its functional groups need to
be chemically modified. In contrast, keratin is a family of fibrous proteins containing cell adhesion RGD sequences similar to fibronectin. In
the form of 3D scaffolds produced by lyophilization method, keratin
supported hADSCs adhesion, proliferation, and differentiation and the
constructs were proved to shorten wound healing time and accelerate
epithelialization in vivo [197]. Gelatin is a biopolymer product of either
partial acid or alkaline hydrolysis of animal collagen. Gelatin scaffolds
showed good affinity to human dermal fibroblasts that made them
suitable for skin regeneration scaffolds [198] while the high viability of
hUVSMCs (human umbilical vein smooth muscle cells) cultivated on
gelatin nanofibre scaffolds indicated a high potential for muscle regeneration [199]. However, disadvantages of natural polymer scaffolds
are weak biomechanics, low stiffness [200], inappropriate degradation
rate with tissue regeneration rate, high price, and lack of surface specificity. Additionally, some natural materials may incite inflammation
in body right [201]. Most natural polymers exhibit good biocompatibility together with poor processability.
proteases), and sustained release of different bioactive moieties immobilized by covalent or non-covalent bonds to the peptide chains.
Moreover, without additional external growth factors, synthetic glycosaminoglycan (GAG) mimetic peptide nanofibre scaffolds were found
to be able to induce neovascularization and cardiomyocyte differentiation for the regeneration of cardiovascular tissue in mice [176].
The GAG-mimic scaffold increased the VEGF expression and localization in the ischemic area and was able to induce activation and migration of cardiac stem cells. Peptide nanofibres were also used to
transplant pancreatic islets with improved both numbers of instrumental vessels and viability of islets in diabetic rats [177]. A glucosaminoglycan – heparin, shows growth factor- (such as VEGF, FGFβ,
TGFβ1) and cell surface receptor-binding affinity. For that reason, it
was used in constructing self-assembling peptide (RAD16-I)-heparin
fibrous hydrogel scaffold that was found to improve cell survival and
differentiation of cells undergoing chondrogenesis probably because of
clustering of receptors in the presence of growth factors associated to
ECM and, therefore, signalling to the receptors [178]. The potential of
these synthetic biomaterials lies in the variability in composition,
peptide length and scaffold design. Because of their low mobility and
high viscosity, self-assembling peptides are mostly used for local protein, gene, drug, etc. delivery [179]. Nonetheless, the control over the
length and structure of supramolecular is a great challenge influencing
their application and biomedical safety. Additionally, it is hard to
consider the effects of different combinations of peptides on binding
ability with other molecules and the motif (amino acid sequence with a
specific function) design assembly on functionalization and expectable
bio-properties of the peptide sequence.
6.2. Natural polymer scaffolds
Natural polymers are biodegradable and bioactive materials that
can be classified into proteins, polysaccharides and polynucleotides
(Table 1). As the main protein in ECM, collagen 3D scaffolds are widely
used in cartilage [180], vasculature [181], nerve [182,183], and muscles [184] regeneration. Without additional modification, collagen
hydrogels display the intrinsic ability of biological recognition such as
receptor-binding ligand presentation and susceptibility to cell-triggered
proteolytic degradation. The basic molecular unit of collagen is tropocollagen containing triple helix structure. It is the most abundant protein (25–35% of the total body content) in ECM of various connective
tissue in the human body [185]. When honeycomb collagen sponge
scaffolds containing bone marrow stromal cells (BMSCs) were implanted into rat model, better sensory and motor recovery of dorsal root
ganglion than only honeycomb scaffold was observed [186].
Chitosan is a polysaccharide compound derived from partial deacetylation of chitin extracted primary of shellfish sources. Self-healing
chitosan-based hydrogel scaffolds with 1.5 kPa stiffness demonstrated
in vivo proliferation and differentiation of neuro-progenitors in the
central nervous system of Zebrafish model [187] which made them also
suitable for neuro-regeneration. To facilitate cellular attachment,
spreading, proliferation, and cytoskeletal organization a successful approach is to incorporate integrin-binding peptide sequence such as RGD
(arginine-glycine-aspartic acid) in natural hydrogel scaffolds [188]. The
pro-inflammatory fibrinogen that is a natural polymer, contains multiple cell-binding motifs including RGD. Zaviscova et al. examined the
neurogenic potential of RGD-modified (20 mg/mL) hydroxyphenyl derivate of HA-based hydrogel (HA-PH-RGD), fibronectin-modified HAPH-RGD (HA-PH-RGD/F), and HA-PH-RGD/F combined with mesenchymal stem cells (MSCs) and found out that after injection of HAPH-RGD and HA-PH-RGD/F in rats the density of neuro-filament fibres
in the leisure increased and neo-vascularization was supported while
these effects were further increased by adding MSCs to HA-PH-RGD/F
[189]. These transplanted cells released trophic factors that provided
immunomodulatory and lasting neurotrophic effect. When combining
chitosan with pro-inflammatory fibrinogen the scaffolds stimulated
6.3. Natural-synthetic polymer blends and composite scaffolds
On one hand, the poor mechanical properties incapable of maintaining the desired 3D shape of scaffolds comprising only of natural
hydrogels is a critical issue. There was found an inverse relationship
between degradation rate and mechanical strength. On the other, synthetic polymers show low cell affinity due to hydrophobicity and lack of
cell recognition sites but they have high flexibility in modification. For
that reason, combinations of natural and synthetic polymers have been
used to design scaffolds with enhanced biodegradability, cell attachment, and hydrophilicity [202]. For example, solid-state synthetic
thermoplastic biopolymers such as PCL and PLGA could be dispensed
together with natural hydrogels with or without different cells seeding.
Shim at al. proposed manufacturing of enhanced 3D bioprinted constructs from PCL and alginate (4% w/v) where PCL acted as mechanically stable framework whereas alginate was responsible for the real
tissue cell arrangement [203]. The dual cell-laden (osteoblasts and
chondrocytes) scaffold enabled viability and proliferation of printed
cells while retaining their initial position which is important for the
regeneration of anatomically complex such as osteochondral tissues. In
another study, chitosan microparticles (10 and 20 wt%) that naturally
promote cellular adhesion without additional functionalization, were
used for blending of the porous scaffolds with 10 and 15 wt% PCL
[204]. 10 wt% PCL with 10 wt% chitosan improved not only storage
and loss moduli but also the in vitro production of type II collagen by
chondrocytes.
Electrospun microfibrous scaffold from collagen/hyaluronic acidbased poly (L-lactide-co-ε-caprolactone)/PLC (9.5/0.5/20 w/w/w) demonstrated higher water uptake and lower Young modulus as opposed
to control PLC scaffold that made it suitable for clinical wound healing
applications [205]. The complex scaffold better supported the adhesion
283
284
[224]
[225]
Chondrocytes
Human adult renal stem cells
Thermally-induced phase separation
Electrospinning
Cartilage
Renal tubular epithelial lineage
[219]
[220]
[221]
[222]
[223]
rCCs (Rabbit corneal cells)
Keratinocytes
Human dermal fibroblasts
ASCs
hESCs
PCL-collagen radially aligned nanofibre scaffolds
Fibroblast loaded collagen-based construct with PCL mesh
TFG-β1 or gentamicin loaded PCL/collagen nanofibres
Devitalized native cartilage with porous PCL scaffolds
Plasma-treated PLLA: PCL (4:1) nanofibrous scaffolds coated
with Matrigel
PLLA, agar, and gelatin scaffolds
PLLA- fibronectin mimetic peptide fibrous scaffolds
Caco-2 (human epithelial cells)
–
Electrospinning
Electrospinning and membrane surface
functionalization
Modified electrospinning
Hybrid extrusion and inkjet process
Electrospinning
Electrospinning
Electrospinning
Intestinal epithelium
Site-specific drug delivery platform with NIR (near-infrared) and pH-triggering
for synergetic photothermal chemotherapy
Cadaveric corneas and amniotic membranes
Human skin
Wound healing
Cartilage
Auditory nerve
[215]
[216]
Liver
Neurons
[213]
[214]
Sweat gland
Cartilage
Extrusion
Co-extrusion
hASCs
SH-SY5Y
Bioprinting allows obtaining pre-determined 3D architectures from
living multi-potent cells or ECM components that can grow for later
implantation. To maintain the viability of the cells, additional growth
and differentiation factors have to be added to the construction so that
all components analogous of living tissue (cells and ECM) to be generated. During printing, the biomaterial behaves like a liquid that could
[208]
[209]
[210]
[50]
[211]
[212]
Extrusion
Freeze-drying
7. Cell-laden scaffolds
Cartilage
Vascular
Cardiac
Cartilage
Cardiac
Liver
Chondrocytes
hMSCs, hUVECs, hNDFs
hUVECs
Chondrocytes
hAVIC
Multicellular primary cell liver
spheroids
Epithelial progenitor cells
Chondrocytes
Inkjet printing
Extrusion
Extrusion
Extrusion
Extrusion
Extrusion
Collagen and fibrinogen scaffolds
Gelatin and fibrinogen scaffolds
Alginate and methacrylated gelatin scaffolds
Nanofibrillated cellulose and alginate scaffolds
Methacrylated hyaluronan and methacrylated gelatin scaffolds
Thiol hyaluronic acid, thiol gelatin, dECM, and PEG-based
crosslinkers in scaffolds
Gelatin, alginate, EGF, and dermal homogenates scaffolds
Alginate, gellan and BioCartilage (micronized human cartilage
particles) scaffolds
Cell-laden collagen core and alginate sheet scaffolds
Heparin sulphate – laminine mimetic peptide amphiphile
nanofibre scaffold
Nanofibrous PET scaffolds coated with collagen
Polypyrrole-coated paclitaxel-loaded PCL fibrous scaffold
Table 3
Examples for scaffolds of polymer blends with soft tissue application.
Cell type
Fabrication
Biomaterial composition
Application
Ref.
of human umbilical vein endothelial cells (hUVECs) and adipose tissuederived mesenchymal stem cells (ADMSCs) and led to 1.6 fold increase
in total vessel length than pure PCL scaffold. Similarly, bilayer electrospun scaffolds of outer gelatin and inner keratin nanofibres
(~160 nm) with random arrangements on PU dressing enhanced fibroblast proliferation in contrast to gelatin mat [206]. The bilayer
scaffold promoted earlier vascularization and better wound healing
capacity. Other complex scaffold composed of hyaluronic acid hydrogel
with multi-tubular conformation modified with anti-Nago receptor antibody and mixed with PLGA microspheres loaded with brain-derived
neurotrophic factor and VEGF displayed very good biocompatibility,
anti-inflammation activity, spinal repair, and enhanced blood vessel
formation when implanted in rat [207]. Table 3 summarizes recent
studies on polymer blends used for various soft tissue applications together with their production technologies and stem cells tested.
When soft tissues require high mechanical activity like cartilage,
heart valves, blood vessels, or dermis, the scaffold ought to have sufficient strength which is important for the effective transfer of mechanical stimuli [226]. As demonstrated previously, incorporation of
bioactive particles in polymer constructs can influence the mechanical
behaviour, degradability and biological performance of composite
scaffolds. Freeze-dried porous collagen (70 wt%)/HAp (30 wt%) scaffolds with average pore diameter of 147 μm and 89% porosity outperformed the mechanical and osteoconductive properties of individual
HAp and collagen scaffolds which make them potential candidates for
the regeneration of damaged cartilage tissue [227]. Similarly, it was
reported that PLGA/nanoHAp scaffolds substantially improved the
cartilage regeneration as compared with pure PLGA scaffold [228].
Nano-GO (1 mg/ml) incorporated into a gelatin-based matrix that after
stereolithography printing photopolymerized into the hierarchical
scaffold with pores of around 200 μm, demonstrated higher compressive modulus and induced higher cell proliferation and chondrogenic
differentiation of hMSCs when compared to those without GO [229].
Although the reinforced scaffolds achieved load-bearing properties similar to those of cartilage tissue, the hydrogel containing composites
may not be sufficiently stable and their degradability rates should be
further examined.
Additionally, by incorporating various substances, scaffold materials can be modified or functionalized not only to increase material for
hosting the cells but also potent anti-inflammatory activity. In this respect, polyurethane (PU) nano-fibrous scaffolds that could be prepared
in both degradable (suitable for skin tissue engineering) and non-degradable (appropriate for wound dressing application) forms were
blended with antimicrobial substances such as silver nanoparticles
[230], copper oxide (showing additional angiogenesis activity) nanoparticles [231], and antibiotics [232]. Simultaneously, these scaffolds
showed low toxicity towards both keratinocytes [230] and fibroblasts
[231] which make them viable candidates for clinical wound healing
application. Moreover, electrospun composite poly(acrylic acid) (PAA)
nanofibre scaffolds incorporating reduced GO (rGO) exhibited controlled near-infrared photo-thermal release of pre-loaded antibiotics
(ampicillin and cefepime) with release rate depending on the applied
radiation [233]. The nanofibre platform demonstrated the ability to be
re-loaded with antibiotics indicating recyclability and stability. The in
vivo studies on S. aureus infected mice showed excellent would healing
capability making the scaffold a suitable platform that could be adapted
for on-demand delivery of different drugs.
[217]
[218]
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It was reported that GF gradients could effectively direct individual cell
migration and control cell alignment [108]. To fabricate complex cytoarchitectures, Graham et al. printed patterned construct of two populations of cells with high-resolution 3D features including layers and
channels under 200 μm width within cubic mm-scaled structure [252].
They used dispensing nozzle ejecting cell-containing hydrogel-based
bioink where the cellular constructs were encapsulated in a thin gel
layer and printed in an oil phase. The obtained relevant tissue density
was 3 × 107 cell mL−1. The lamellar constructs of both cell types,
human embryonic kidney cell (hEKC) derivate and ovine mesenchymal
stem cell (oMSCs) in homogeneous distribution remained viable and
responsive to growth factors for osteoblasts and chondrocytes differentiation.
For the successful implantation of the biomaterial, biomimetic
structure seeded with cells either donated from the host or from compatible tissue bank is required. Even decellularized ECM (dECM) was
found to be tissue-specific providing crucial cues for cell survival and
function. Consequently, dECM bioinks from adipose (adECM) and cartilage (cdECM) tissue with encapsulated cells were used for the production of 3D porous scaffolds of dECM and PCL framework by multihead tissue/organ building system [240]. The cell viability of the
printed construct was over 95%. Using encapsulated human adipose-
bring different biomimetic architectures. There are three main approaches for developing cell-laden scaffolds (Fig. 6):
a) scaffold-based – this top-down strategy uses bioactive scaffolds
acting like mechanical support for immobilization of cells and
bioactive molecules; although applying good control over the scaffold characteristics (shape and size, biofunctionalization), usually,
the initial cell density, tissue assembly and vascularization are poor;
b) scaffold-free - bottom-up approach where cell aggregates, sheets or
spheroids are used as building blocks of the scaffolds; however, low
biomolecule-functionalization ability and insufficient mechanical
stability are the main disadvantages of this strategy;
c) synergetic or bioassembly – it uses bottom-up random assembly of
multiple functional units (cell-laden modules) and combines the
advantages of both previous strategies.
Such cell-containing microunits well replicate the physiological
conditions (with efficient nutrient and oxygen transfer), possess high
surface-to-volume ratio and can be arranged in the desired geometry.
Nevertheless, this approach suffers from an inability to fabricate complex scaffolds with controlled micrometre-scaled features and when
heterogeneous cell populations are assembled, poor functionality could
occur [234].
As opposed to bioassembly, bioprinting bottom-up technologies use
smaller fabrication units such as cells, cell aggregates or biochemicals
to form scaffolds with ordinate assembly. The cells or bioactive moieties
are included in bioinks of hydrogel precursors or ECM-containing solutions that are printed continuously or in the form of small droplets.
The bioprinting technologies include extrusion-based [235], dropletbased ejection [236], laser-assisted [237], and hybrid methodologies.
Extrusion-based bioprinters use either continuous cell-laden hydrogel
(from chitosan, gelatin, fibrinogen, etc.) filaments or cell spheroids
[238] that could be deposited without a scaffold. When a microtissuebased approach is used, smaller 3D spheroid aggregates in degradable
carrying medium with uniform size and shape are incorporated into
bioink that after printing consolidates into macrotissue [239]. This
technique was successfully used for producing of simple tissues such as
adipose tissue [240], muscles, bones, cartilage [241], 3D vasculature
[235], neural mini-tissues differentiated to functional neurons and
supporting neuroglia [242], and vascularized heart tissue [243]. Extrusion based printers are suitable for the production of large and
complex (branched or tubular) structures but can be deficient for
printing high-resolution or multiple cell-type structures. The dropletbased bioprinters dispense cell-loaded droplets from a nozzle by sonic,
pneumatic or thermal actuation [244]. In valve-based droplet ejection,
the size and number of cells in a droplet can be controlled by valve
opening time or actuation frequency while in acoustic-based, picoliter
cell-encapsulating droplet is generated with a few microseconds pulse
duration. The droplet-based method proposes high-resolution, automation, versatility, and cell viability for the account of clogging and
lower cell density [245]. Bone, cartilage tissue scaffolds [208], microvasculature constructions [246], and cardiac patch [247] have been
fabricated by this method. Laser-assisted bioprinting (LAB) uses laser
irradiation to assemble cell-containing microdroplets ejected from a
ribbon on a substrate producing 2D and sometimes 3D architectures
[248]. Although Koch et al. reported no or below detection deleterious
effect of LAB on printed cells [249], there is a need for evaluation cell
function and damage during and after printing in various hydrogels,
temperature and rheological ranges compatible with bioprinting technologies.
Despite the recent developments of bioprinting, it is still a challenge
to build high-resolution constructs from multiple cell-types patterned
necessary for cell-cell interaction and function. One option is to print
spatially defined gradients of immobilized GFs (at pico- to nanogram
quantities) that influence the differentiation of seeded multi-potential
cells such as neural stem cells [250] or mesenchymal stem cells [251].
Fig. 6. Strategies for tissue regeneration by using cell-laden scaffolds: The topdown approach uses scaffolds biofunctionalized with cells and other biomolecules. The tissue is regenerated after cell proliferation and scaffold degradation.
The bottom-up approach involves cell aggregates, sheets, modules or bioprinted
elements to produce blocks for assembling the scaffolds needed for tissue regeneration.
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derived stem cells (hADSCs) and human inferior turbinate-tissue derived mesenchymal stromal cells (hTMSCs) in adECM and cdECM, respectively, the authors generated a synergetic effect of combining engineered porosity with an increased commitment of stem cells towards
specific differentiation [240].
Although expecting lower liver-specific functions, Cheng et al.
[253] synthesized core-shell microgel scaffolds containing fibroblast
loaded alginate shell and hepatocyte-filled aqueous core. In contrast to
monotypic (without fibroblasts) counterpart, the co-cultured spheroids
demonstrated increased urea synthesis and albumin secretion that are
biomarkers used when the liver is drug-screened. The same research
group was able to co-encapsulate hepatocytes and endothelial progenitor cells (EPCs) into alginate-collagen microgels that showed enhanced hepatocellular functions only at co-culture ratio 5:1, respectively [254]. The increase in EPCs fraction reduced the performance of
hepatocytes. Moreover, it was hard to supply sufficient growth factors
to the encapsulated in hydrogel cells for their differentiation.
In another study, Min et al. printed complex full-thickness skin
construct containing pigmentation [255]. For that purpose, the scaffold
consisted of five layers of collagen-precursor solution bioprinted in
small cell-containing droplets and crosslinked through neutralization
with NaHCO3. A suspension of fibroblasts was embedded in 2–4 layers.
On the top of the dermal layer, suspension of melanocytes was printed,
followed by primary keratinocyte layers above it. When submerged in
air-liquid interface (ALI) culture, the formation of distinct dermal and
epidermal layers with terminal differentiation of keratinocytes was
observed. The melanocytes and keratinocytes in the epidermal layer
formed dark-pigmented non-uniform clusters underlining the need for
optimization the pigmentation homogeneity.
It follows that high resolution, high-speed organization and high
cell viability are key requirements for cell-laden 3D scaffold fabrication.
After that, efficient vascularization precisely placed at a high level of
density cells is an important precondition for successful implantation.
Recently, using multiple bioinks loaded with vein endothelial cells,
mesenchymal stem cells, and dermal fibroblasts for the synthesis of
ECM, Kolesky et al. [209] bioprinted 3D vascularized tissue with
thickness over 1 cm that actively perfused GFs to differentiate hMSCs
towards osteogenesis. The 3D constructs underwent fast tissue maturation where the scaffold provided controllable growth, maturation,
and proper differentiation. Similarly, Mori et al. printed skin equivalent
containing perfusable vascular channels coated with endothelial cells
[256]. The vascular channels with conventional epidermal and dermal
morphology were opened in the perfused skin-equivalent because of the
pressure generated by perfusion. The results confirmed the barrier
function of the epidermal layer that could be controlled by changing the
concentration of normal human epidermal keratinocytes (NHEKs). The
vascular channel permeability was size-selective and controlled by
VEGF. These successive constructions of vascular channels could make
the implementation of sweat glands, hair, nervous and immune system
components possible. However, still pending problems are the poor
mechanical properties and limited biological interactions of hydrogels
that could be effectively replaced by alternative types of bioinks such as
supramolecular hydrogels that propose an easy modification, self-assembly properties, and tunable bioactive behaviour [257].
the biological effects of the complex scaffolds usually at the expense of
applying complex engineered technologies. The advanced composite
scaffolds could also segregate in two or more layers with smooth discrete gradients in chemical or physical properties that are able to organize and guide, and maintain multiple cellular phenotype structures
and heterogeneous extracellular matrix morphogenesis.
Although considerable progress has been made in 3D scaffold research, many challenged have to be resolved that make the choice of a
suitable scaffold and biomaterial quite difficult. It should be taken into
account that cell- and growth factor-incorporated scaffolds are difficult
to be stored because of their low viability and stability. Moreover, the
long term release of immobilized substances does not always lead to
achieving the desired effects in the microenvironment promoting
functional recovery or providing trophic or anti-inflammatory support.
Some of the major challenges of 3D scaffold involve:
a. Tuning the structural, biomechanical properties, and degradation
rate of scaffolds with optimized surface characteristics to enhance
cell contacts and ECM deposition depending on the intended application;
b. Efforts in combining different biomaterials with incorporated
bioactive molecules and different fabrication methods for producing
superior scaffolds for a specific application without burst release and
easy inactivation;
c. Increased neurovascularization that inhibits localized necrosis and
implant failure;
d. Formation of multi-tissue types scaffolds capable of spatial and
functional cell regulation;
e. Increased resolution and accuracy of scaffold producing technologies enabling precise replication of fine tissue composition;
f. Minimizing the adverse effects and secondary damages;
g. Simplifying the process of fabrication;
h. Mass production outside the laboratory environment.
Resolving these issues, the 3D scaffold development may lead to
improved tissue regeneration and move the application of 3D scaffold
from screening purpose and evaluation in animals to intended clinical
use. There is still a combination of factors including space constraint,
biomechanical comparability, and other ultimately effecting changes in
cell microenvironment predisposing cells toward attachment and influencing the regeneration process that is presently unknown.
Declaration of competing interest
None.
Acknowledgement
This study was funded by the National Science Fund of Bulgaria
(NSFB), Contract № DN 07/3 (2016), Gradient functional nanocoatings
produced by vacuum technologies for biomedical applications.
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