An Introduction to 3D Bioprinting: Possibilities, Challenges and Future Aspects

Abstract

Bioprinting is an emerging field in regenerative medicine. Producing cell-laden, three-dimensional structures to mimic bodily tissues has an important role not only in tissue engineering, but also in drug delivery and cancer studies. Bioprinting can provide patient-specific spatial geometry, controlled microstructures and the positioning of different cell types for the fabrication of tissue engineering scaffolds. In this brief review, the different fabrication techniques: laser-based, extrusion-based and inkjet-based bioprinting, are defined, elaborated and compared. Advantages and challenges of each technique are addressed as well as the current research status of each technique towards various tissue types. Nozzle-based techniques, like inkjet and extrusion printing, and laser-based techniques, like stereolithography and laser-assisted bioprinting, are all capable of producing successful bioprinted scaffolds. These four techniques were found to have diverse effects on cell viability, resolution and print fidelity. Additionally, the choice of materials and their concentrations were also found to impact the printing characteristics. Each technique has demonstrated individual advantages and disadvantages with more recent research conduct involving multiple techniques to combine the advantages of each technique.

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materials
Review
An Introduction to 3D Bioprinting: Possibilities,
Challenges and Future Aspects
Željka P. Kaˇcarevi´c 1,* , Patrick M. Rider 2, Said Alkildani 3, Sujith Retnasingh 4,
Ralf Smeets 5,6, Ole Jung 6, Zrinka Ivaniševi´c 7and Mike Barbeck 2,5,8
1Department of Anatomy Histology, Embryology, Pathology Anatomy and Pathology Histology,
Faculty of Dental Medicine and Health, University of Osijek, 31000 Osijek, Croatia
2Botiss Biomaterials, Hauptstraße 28, 15806 Zossen, Germany; patrick.rider@botiss.com (P.M.R.);
mike.barbeck@icloud.com (M.B.)
3
Department of Biomedical Engineering, Faculty of Applied Medical Sciences, German-Jordanian University,
11180 Amman, Jordan; saidkildani@gmail.com
4Institute for Environmental Toxicology, Martin-Luther-Universität, Halle-Wittenberg and Faculty of
Biomedical Engineering, Anhalt University of Applied Science, 06366 Köthen, Germany;
sujiroshi@gmail.com
5Department of Oral and Maxillofacial Surgery, University Hospital Hamburg-Eppendorf,
20246 Hamburg, Germany; r.smeets@uke.de
6
Department of Oral Maxillofacial Surgery, Division of Regenerative Orofacial Medicine, University Medical
Center Hamburg-Eppendorf, 20246 Hamburg, Germany; ol.jung@uke.de
7Department of Dental Medicine, Faculty of Dental Medicine and Health, University of Osijek,
31000 Osijek, Croatia; zrinkaivan@gmail.com
8BerlinAnalytix GmbH, 12109 Berlin, Germany
*Correspondence: zeljkapericc@gmail.com; Tel.: +49-3020-6073-9858
Received: 1 October 2018; Accepted: 2 November 2018; Published: 6 November 2018


Abstract:
Bioprinting is an emerging field in regenerative medicine. Producing cell-laden,
three-dimensional structures to mimic bodily tissues has an important role not only in tissue
engineering, but also in drug delivery and cancer studies. Bioprinting can provide patient-specific
spatial geometry, controlled microstructures and the positioning of different cell types for the
fabrication of tissue engineering scaffolds. In this brief review, the different fabrication techniques:
laser-based, extrusion-based and inkjet-based bioprinting, are defined, elaborated and compared.
Advantages and challenges of each technique are addressed as well as the current research status
of each technique towards various tissue types. Nozzle-based techniques, like inkjet and extrusion
printing, and laser-based techniques, like stereolithography and laser-assisted bioprinting, are all
capable of producing successful bioprinted scaffolds. These four techniques were found to have
diverse effects on cell viability, resolution and print fidelity. Additionally, the choice of materials
and their concentrations were also found to impact the printing characteristics. Each technique has
demonstrated individual advantages and disadvantages with more recent research conduct involving
multiple techniques to combine the advantages of each technique.
Keywords:
additive manufacturing; 3D scaffolds; inkjet; extrusion; stereolithography; laser-assisted;
rapid prototyping
1. Introduction
Bioprinting is a subcategory of additive manufacturing (AM), also known as three-dimensional
(3D) printing. It is defined as the printing of structures using viable cells, biomaterials and biological
molecules [
1
,
2
]. Bioprinting must produce scaffolds with a suitable microarchitecture to provide
Materials 2018,11, 2199; doi:10.3390/ma11112199 www.mdpi.com/journal/materials

Materials 2018,11, 2199 2 of 21
mechanical stability and promote cell ingrowth whilst also considering the impact of manufacture on
cell viability; for instance, chemical cytotoxicity caused by the use of solvents or pressure-induced
apoptotic effect produced during the extrusion of material. A significant benefit of bioprinting is
that it prevents homogeneity issues that accompany post-fabrication cell seeding, as cell placement is
included during fabrication.
The advantage of homogeneously distributed cell-laden scaffolds has been demonstrated by faster
integration with the host tissue, lower risk of rejection and most importantly, uniform tissue growth
in vivo
[
3
–
6
]. Conventional cell seeding techniques are either static or dynamic, and while the latter
one results in better seeding efficiency and cell penetration into the scaffold, it is known affect cell
morphology [7].
Immediate vascularization of the implanted scaffolds is highly critical [
8
,
9
]. With proper
vascularization, the scaffolds are provided with an influx of oxygen/nutrients and an efflux of carbon
dioxide/by-products; preventing core necrosis. Vascularization also supports the implants with
remodelling [
10
]. Bioprinting techniques have been employed to fabricate microvascular-like structures
and have the potential to position endothelial cells within the 3D structures as a prevascularization
step prior to implantation [11].
Bioprinting can be applied in a clinical setting, where it can be used to create regenerative scaffolds
to suit patient specific requirements [
12
]. The process of applying bioprinting to a clinical setting is
depicted in Figure 1. To begin with, imaging modalities such as CT, MRI and ultrasound can be used
to create a digital 3D model of the tissue defect. Using computer aided design (CAD), the internal and
external architecture of the scaffold, such as porosity and pore sizes, can be incorporated into the 3D
model of the tissue defect. In consideration of the defect type, location and requirements, a selection of
materials, cell types and bioactive molecules, can be used to fabricate a bioink for printing. Cell laden
structures are then manufactured using bioprinting technology and are then placed either in cell
culture or directly implanted into the patient.
Materials 2018, 11, x FOR PEER REVIEW 2 of 22
on cell viability; for instance, chemical cytotoxicity caused by the use of solvents or pressure-induced
apoptotic effect produced during the extrusion of material. A significant benefit of bioprinting is that
it prevents homogeneity issues that accompany post-fabrication cell seeding, as cell placement is
included during fabrication.
The advantage of homogeneously distributed cell-laden scaffolds has been demonstrated by
faster integration with the host tissue, lower risk of rejection and most importantly, uniform tissue
growth in vivo [3–6]. Conventional cell seeding techniques are either static or dynamic, and while
the latter one results in better seeding efficiency and cell penetration into the scaffold, it is known
affect cell morphology [7].
Immediate vascularization of the implanted scaffolds is highly critical [8,9]. With proper
vascularization, the scaffolds are provided with an influx of oxygen/nutrients and an efflux of
carbon dioxide/by-products; preventing core necrosis. Vascularization also supports the implants
with remodelling [10]. Bioprinting techniques have been employed to fabricate microvascular-like
structures and have the potential to position endothelial cells within the 3D structures as a
prevascularization step prior to implantation [11].
Bioprinting can be applied in a clinical setting, where it can be used to create regenerative
scaffolds to suit patient specific requirements [12]. The process of applying bioprinting to a clinical
setting is depicted in Figure 1. To begin with, imaging modalities such as CT, MRI and ultrasound
can be used to create a digital 3D model of the tissue defect. Using computer aided design (CAD),
the internal and external architecture of the scaffold, such as porosity and pore sizes, can be
incorporated into the 3D model of the tissue defect. In consideration of the defect type, location and
requirements, a selection of materials, cell types and bioactive molecules, can be used to fabricate a
bioink for printing. Cell laden structures are then manufactured using bioprinting technology and
are then placed either in cell culture or directly implanted into the patient.
Figure 1. Schematic of Bioprinting Scaffolds for clinical use. Digital 3D images obtained from CT,
MRI or ultrasound, are used to design a suitable scaffold with 3D slicing and CAD software;
materials from printing are chosen depending upon the application, and can consist of polymers,
ceramics, and bioactive components; cells are selected dependent on the application, a bioink can
consist of singular or multiple cell types; post-fabrication 3D culture can be used for characterization,
assessment and ultimately implantation. 3D printing is both time and cost effective, enabling fast
adjustments and implementation of designs [13]. Designs can be made to match exact defect
geometries, improving the union between implant and native tissue, thereby enhancing tissue
integration [14]. Additive manufactured scaffolds have shown satisfactory accuracy matching the
designs [15–17]. Different types of tissues and organs have been produced using bioprinting, for
instance; blood vessels [18], heart tissue [19], skin [20,21], liver tissue [5], neural tissue [22], cartilage
[23] and bone [24].
Figure 1.
Schematic of Bioprinting Scaffolds for clinical use. Digital 3D images obtained from
CT, MRI or ultrasound, are used to design a suitable scaffold with 3D slicing and CAD software;
materials from printing are chosen depending upon the application, and can consist of polymers,
ceramics, and bioactive components; cells are selected dependent on the application, a bioink can
consist of singular or multiple cell types; post-fabrication 3D culture can be used for characterization,
assessment and ultimately implantation. 3D printing is both time and cost effective, enabling fast
adjustments and implementation of designs [
13
]. Designs can be made to match exact defect geometries,
improving the union between implant and native tissue, thereby enhancing tissue integration [
14
].
Additive manufactured scaffolds have shown satisfactory accuracy matching the designs [
15
–
17
].
Different types of tissues and organs have been produced using bioprinting, for instance; blood
vessels [
18
], heart tissue [
19
], skin [
20
,
21
], liver tissue [
5
], neural tissue [
22
], cartilage [
23
] and bone [
24
].

Materials 2018,11, 2199 3 of 21
The ultimate aim of bioprinting is to provide an alternative to autologous and allogeneic tissue
implants, as well as to replace animal testing for the study of disease and development of treatments.
In this review, the main bioprinting techniques are discussed: inkjet-based, extrusion-based and
laser-assisted, including their basic mechanisms and current challenges. Tables 1–3provide an
overview of recent research for each technique.
An important component of bioprinting is the use of bioinks. Bioinks consist of biomaterials that
can be used to encapsulate cells and incorporate biomolecules. Cell laden bioinks are hydrogel-based,
as hydrogels have a high water content that is beneficial for cell survivability and shielding the cells
from fabrication induced forces. The main properties of a bioink that need to be considered before
printing include its viscosity, gelation and crosslinking capabilities. These properties can significantly
affect print fidelity (construct stability and print deviation from the computer aided designs) as well
as cell viability, proliferation and morphology after printing [
25
]. To produce a hydrogel that can
both support and protect the cells, whilst at the same time provide a structurally secure scaffold is
challenging, as these characteristics have different mechanical requirements. Stiff hydrogels have
denser networks that might put the cells under pressure during encapsulation, as well as hinder their
migration [
26
]. Ultimately, the hydrogel properties need to be balanced between structural fidelity and
cell suspension.

Materials 2018,11, 2199 4 of 21
Table 1.
Recent
in vitro
studies. AG—Agarose, SA—Sodium alginate, PLA—Polylactide fibers, GelMA—gelatin methacryloyl, HUVECs—Human umbilical vein
endothelial cells, PEGDA—poly(ethylene glycol) diacrylate, ATCC—Mouse neural stem cell lines, BrCa—breast cancer cells, MSCs—marrow mesenchymal stem cells,
Nha—nanocrystalline hydroxyapatite.
Biomaterials Cells Results Significance Reference
Extrusion-based techniques
SA SA/collagen SA/AG Chondrocytes
Printed SA/collagen scaffold in cell culture
showed enhanced cell proliferation, cartilage
specific gene expression and cell adhesion.
SA/collagen is a potential bioink base
material for cartilage regeneration Yang et al., 2017 [27]
Alginate PLA fibers Human chondrocytes Printed cells showed a high cell viability (80%).
The addition of sub-micron PLA fibers can
be used to improve hydrogel mechanical
properties
Kosik-Kozioł 2017 [28]
GelMA HUVECs
Printed cells form lumen- like structure of the
endothelium and contracted with an
approximate rate of 60 bpm for up to 7–10 days
when cultured.
Successfully demonstrated the 3D printing
of endothelialized-myocardium-on-a chip. Zhang 2016 [19]
Laser-assisted bioprinting
Human Osseous Cell Sheets HUVECs
Printed cell exhibits the formation of tubule-like
structures within the biopaper after 21 days of
culture.
Demonstration of self-assembled cell
sheets for the soft tissue regeneration. Kawecki 2018 [29]
Stereolithography
PEGDA and GelMA
MCF-7 breast cancer cell, HUVECs,
C2C12 skeletal muscle cells,
osteoblasts, fibroblasts,
mesenchymal cells.
Fabricated structure exhibited high cell viability,
proliferation and metabolic activity.
Demonstrated the flexibility of
stereolithography for printing different
cell types
Miri 2018 [26]
GelMA and graphene
nanoplatelets ATCC
The printed cells had differentiated, produced
well-defined architectures and homogenous cell
distribution.
Successfully demonstrated the printing
neural stem cells Zhu 2016 [30]
GelMA and nHA BrCa and MSCs Printed MSCs secreted macromolecules that
promoted BrCa growth.
Successful model for the investigation of
post-metastatic breast cancer progression
in bone.
Zhou 2016 [31]
Inkjet-based techniques
Cell suspension Porcine Schwann cells, Neuronal
analogue NG108-15 cells
Printed neuronal cells exhibited high cell
viabilities as well as earlier and longer neurite
growth than unprinted cells.
Can be incorporated into large tissue
models to include an established neuronal
network before implantation.
Tse 2016 [32]
Alginate
Primary feline adult
cardiomyocytes, HL1 cardiac
muscle cell line
Cells remained viable in a large scaffold.
Scaffold pulsated under electrical stimulation. Successfully printed myogenic tissue Xu 2009 [33]

Materials 2018,11, 2199 5 of 21
Table 2.
Recent
in vivo
studies. Abbreviations: PU—poly(urethane), PCL—poly(caprolactone), hASCs—human adipose-derived stem cells, NSCs—neural stem
cells, PEG—poly(ethylene glycol), HUVECs—human umbilical vein endothelial cells, iPSCs—induced pluripotent stem cells, CM—cardiomyocytes, bMSCs—bone
marrow-derived mesenchymal stem cells, ROB—rat osteoblasts, TCP—tricalcium phosphates, HMECs—human microvascular endothelial cells.
Biomaterials Cells Results Significance Reference
Extrusion-based techniques
Hyaluronic acid, Gelatin,
Glycerol, Fibrinogen, PU
Human fibroblasts, Human
keratinocytes
Subcutaneous implants in rats reduced wound
area to <40% after 14 days. Regenerated skin
tissue consisted of epidermis and dermis layers
Novel method to fabricate patient-specific
tissue construct to reconstruct facial
skin wounds
Seol, 2018 [34]
Human decellularized adipose
tissue, PCL hASCs The scaffolds proved to be adipo-inductive and
exhibited adequate tissue infiltration
Demonstration of a clinically viable method
of soft tissue regeneration Pati, 2015 [35]
PU nanoparticles NSCs Implanted in adult zebrafish repaired traumatic
brain injuries and restored function
3D printing system that does not involve the
use of heat, toxic organic solvents, toxic
photoinitiators or UV for crosslinking
Hsieh, 2015 [36]
Alginate/gelatin,
Alginate/hyaluronic acid,
Alginate/Matrigel
INS1E-ß cells, Islets, (human
and mouse)
Implanted subcutaneously in mice, exhibited
metabolic activity after 7 days
Demonstrates possibility of encapsulating
and printing human islets for islet
transplantation applications
Yanez, 2015 [37]
Alginate, Fibrinogen, PEG HUVECs, iPSCs-derived CMs
Subcutaneous implants in NOD-SCID mice
developed a vascular network and CMs
exhibited maturation after 2 weeks
Demonstrates an advantageous printing
design where extruded filament was
composed of 2 different inks
Maiullari, 2018 [38]
PCL, Sodium alginate
Rabbit bMSCs, Rabbit
chondrogenic bMSCs, Rabbit
respiratory endothelial cells
Neocartilage and neovascularization in rabbits
after 12 weeks of tracheal implantation
Demonstrates fabrication of an artificial
trachea with two cell types via additive
manufacturing
Bae, 2018 [39]
PEG, Laponite XLG, Hyaluronic
acid ROBs Implanted into rat tibias, exhibited new bone
formation after 12 weeks
Demonstrates benefit of extruding the
scaffold support material and bioink
separately, however combined into one
printing process
Xinyun Zhai, 2018 [40]
PCL/TCP/Pluronic®F127,
PCL/Pluronic®F127
Human amniotic-derived stem
cells, Rabbit ear chondrocytes,
Rabbit myoblasts
Implanted into rats, scaffolds with different cell
types produced: newly formed vascularized
bone tissue; vasculature with physiologically
relevant mechanical properties; nerve integration
Showed significant improvements compared
to acellular scaffolds for myogenic and
osteogenic tissues
Kang, 2016 [41]
Laser-based techniques
Collagen Mouse fibroblasts, Human
keratinocytes
Subcutaneous implants in nude mice form
multi-layered epidermis and vascularization
towards the printed cells, after 11 days
Utilization of a laser-assisted printing
process in adding cells to commercially
available skin grafts
Michael, 2013 [21]
Inkjet-based techniques
Fibrin HMECs Printed cells form confluent tubular structure
after 21 days
Promising approach for human
microvascular tissue engineering Cui, 2009 [42]
Collagen, Thrombin, Fibrinogen
Neonatal human dermal
fibroblasts and epidermal
keratinocytes, Dermal
microvascular endothelial cells
Printed scaffolds exhibited 17% better wound
contraction after 6 weeks in nude mice
Positioning of microvascular endothelial cells
on fibroblast/keratinocyte grafts seemed to
be advantageous over commercially
available fibroblast/keratinocyte grafts
Marchioli, 2015 [43]

Materials 2018,11, 2199 6 of 21
Table 3.
Recent in situ studies. Abbreviations: IPFP—Human infrapatellar fat pad-derived adipose stem cells, GelMA—gelatin methacryloyl, HAMa—hyaluronic
acid–methacrylate hydrogel, PEGDMA—Poly(ethylene glycol) dimethacrylate, AFS—Amniotic fluid-derived stem cells, MSCs—bone marrow-derived mesenchymal
stem cells.
Biomaterials Cells Results Significance Reference
Extrusion-based techniques
HA-GelMA MSCs Demonstrated cultured cells directly into
the cartilage defect in sheep.
Directly reconstruction of cartilage using
extrusion printing. Di Bella 2017 [44]
Laser-based techniques
nHA MSCs
Printed cells exhibits the presence of
pulsating blood vessels after bone defect
achievement.
Scaffold was successfully printed in the
mouse calvaria defect model in vivo. Keriquel 2010 [45]
Inkjet-based techniques
PEGDMA Human chondrocytes
Printed directly onto the femoral condyles
defects showed enhanced tissue
integration.
Improved integration by direct in situ
printing. Cui 2012 [46]
Fibrinogen-collagen AFS and MSCs
Used to repair full thickness wounds in the
backs of mice, histological test shows the
presence of blood vessel in the
subcutaneous adipose tissue.
Potential to quickly close full thickness
burns and enable revascularization of the
tissue.
Skardal 2012 [47]

Materials 2018,11, 2199 7 of 21
2. Inkjet-Based Bioprinting
First attempts to print live cells was performed using a specially adapted commercially available
inkjet printers [
1
]. An initial problem encountered when developing inkjet bioprinting was that the
cells died during printing due to instantaneous drying out once on the substrate. The problem was
overcome by encapsulating the cells in a highly hydrated polymer, this led to the development of
cell-loaded hydrogels [
48
]. Inkjet bioprinting allows for the precise positioning of cells, with some
studies achieving as few as a singular cell per printed droplet [
49
]. Cells and biomaterials are patterned
into a desired pattern using droplets, ejected via thermal or piezoelectric processes, depicted in
Figure 2[1,50].
Materials 2018, 11, x FOR PEER REVIEW 7 of 22
2. Inkjet-Based Bioprinting
First attempts to print live cells was performed using a specially adapted commercially
available inkjet printers [1]. An initial problem encountered when developing inkjet bioprinting was
that the cells died during printing due to instantaneous drying out once on the substrate. The
problem was overcome by encapsulating the cells in a highly hydrated polymer, this led to the
development of cell-loaded hydrogels [48]. Inkjet bioprinting allows for the precise positioning of
cells, with some studies achieving as few as a singular cell per printed droplet [49]. Cells and
biomaterials are patterned into a desired pattern using droplets, ejected via thermal or piezoelectric
processes, depicted in Figure 2 [1,50].
Thermal-based inkjet printing uses a heated element to nucleate a bubble. The bubble causes a
build-up pressure within the printhead, which leads to the expulsion of a droplet. The thermal
element can reach temperatures between 100 °C to 300 °C. Initially there have been concerns that
such high temperatures would damage the cells [51], however research has shown that the high
temperatures are localized and are only present for a short time span [11,52].
Piezoelectric-based apparatus uses acoustic waves to eject the bioink. This mechanism limits the
use of highly concentrated and viscous bioinks as their viscosity dampens the applied
acoustic/pressure waves, hindering the ejection of a droplet [53]. A low viscosity is achieved by
using low concentration solutions, a limiting factor for producing 3D structures [50].
Inkjet printing offers a high resolution of up to 50 µm [54]. Most inkjet bioprinters provide a
high cell viability, and although there is the potential for induced sheer stresses to damage the cells,
most research indicates that this is not the case [55,56]. The advantages of inkjet-based bioprinting
include high print speeds, low cost and a wide availability, however problems include low droplet
directionality and unreliable cell encapsulation due to the low concentration of the ink [1].
Figure 2. Schematic of Inkjet-based Bioprinting. Thermal inkjet uses heat-induced bubble nucleation
that propels the bioink through the micro-nozzle. Piezoelectric actuator produces acoustic waves that
propel the bioink through the micro-nozzle.
Cui et al. developed a 3D printed bone-like tissue using poly (ethylene glycol) dimethacrylate
(PEGDMA), that had a similar compressive modulus to natural bone, and bioceramic nanoparticles
[57]. Human mesenchymal stem cells (hMSCs), PEGDMA with hydroxyapatite (HA) and/or bioglass
(BG) nanoparticles were bioprinted into bone tissue scaffolds. The bioceramic nanoparticles were
used to mimic the native bone tissue microenvironment and stimulated the differentiation of stem
Figure 2.
Schematic of Inkjet-based Bioprinting. Thermal inkjet uses heat-induced bubble nucleation
that propels the bioink through the micro-nozzle. Piezoelectric actuator produces acoustic waves that
propel the bioink through the micro-nozzle.
Thermal-based inkjet printing uses a heated element to nucleate a bubble. The bubble causes
a build-up pressure within the printhead, which leads to the expulsion of a droplet. The thermal
element can reach temperatures between 100
◦
C to 300
◦
C. Initially there have been concerns that
such high temperatures would damage the cells [
51
], however research has shown that the high
temperatures are localized and are only present for a short time span [11,52].
Piezoelectric-based apparatus uses acoustic waves to eject the bioink. This mechanism
limits the use of highly concentrated and viscous bioinks as their viscosity dampens the applied
acoustic/pressure waves, hindering the ejection of a droplet [
53
]. A low viscosity is achieved by using
low concentration solutions, a limiting factor for producing 3D structures [50].
Inkjet printing offers a high resolution of up to 50
µ
m [
54
]. Most inkjet bioprinters provide
a high cell viability, and although there is the potential for induced sheer stresses to damage the cells,
most research indicates that this is not the case [
55
,
56
]. The advantages of inkjet-based bioprinting
include high print speeds, low cost and a wide availability, however problems include low droplet
directionality and unreliable cell encapsulation due to the low concentration of the ink [1].
Cui et al. developed a 3D printed bone-like tissue using poly (ethylene glycol) dimethacrylate
(PEGDMA), that had a similar compressive modulus to natural bone, and bioceramic nanoparticles [
57
].
Human mesenchymal stem cells (hMSCs), PEGDMA with hydroxyapatite (HA) and/or bioglass (BG)
nanoparticles were bioprinted into bone tissue scaffolds. The bioceramic nanoparticles were used

Materials 2018,11, 2199 8 of 21
to mimic the native bone tissue microenvironment and stimulated the differentiation of stem cells
towards osteogenic linage. There was significant difference between compressive mechanical strengths
of pure PEG and PEG-HA scaffolds (~0.35 MPa); however, mechanical strength dropped significantly
for PEG-BG scaffolds. Incubation of scaffolds in cell culture for 21 days seemed to increase modulus in
all samples except for PEG-BG. The interaction of hMSCs and HA nanoparticles produced highest cell
viability of 86% compared to the other scaffolds.
Inkjet bioprinting has demonstrated excellent cell viabilities and the potential for creating a neural
network in printed organs. Tse et al. fabricated neural tissue by bioprinting porcine Schwann cells and
neuronal NG 108-15 cells using a piezoelectric inkjet printer [
32
]. Neuronal and glial cell viabilities of
86% and 90% were observed immediately after printing. Proliferation rate of the printed cells was close
to those which weren’t printed. The printed cells seemed to have developed neurites that elongated
after 7 days.
Cardiac tissue with a beating cell response was engineered by Aho et al. using feline
cardiomyocytes HI.1 cardiac muscle cells and an alginate hydrogel. The tissue was fabricated
by printing layers of CaCl
2
into an alginate hydrogel precursor solution to facilitate crosslinking.
The results suggested that cardiac cells attached to the alginate, effectively mimicked the native cardiac
ECM. The printed cardiac tissues exhibited contractile properties under mild electrical stimuli [33].
Min et al. fabricated full thickness skin models with pigmentation using an inkjet technique [
58
].
Dermal models was fabricated from fibroblast-laden collagen. After culturing for 1 day in fibroblast
medium, keratinocytes were printed on top of the dermal model and put in culture for another day.
Melanocytes were then printed onto the model and further cultured in melanocyte medium for 2 h.
The entire model was subjected to air-liquid-interface for 4 days. The construct had distinctive
epidermal and dermal layers. Keratinocytes reached maturation and melanocytes resulted in
freckle-like pigmentation (without chemical or UV stimuli). Sodium carbonate was used for
crosslinking. Yanez et al. investigated the wound healing capabilities of bioprinted skin grafts [
59
].
Skin grafts were fabricated by printing fibrinogen solution onto to a layer of collagen that was laden
with human dermal fibroblasts (NHDFs). A subsequent layer of thrombin, laden with human dermal
microvascular endothelial cells (HMVECs) was bioprinted onto the fibrinogen. Finally, collagen
laden with neonatal human epidermal keratinocytes (NHEKs) was printed onto the fibrin-HMVEC
layer. The grafts were incubated for 24 h and transplanted subcutaneously in to the backs of mice.
Wounds treated with the bioprinted scaffold had completely healed after 14–16 days, whereas wounds
treated without the graft healed in 21 days.
Inkjet bioprinting is of great interest as it exhibits high resolution and cell viability. With this
process, accurate position of multiple cell types is possible [
49
,
60
]. However, the limitations of vertical
printing and restricted viscosities may mean that inkjet bioprinting needs to be combined with other
printing techniques for future developments.
3. Laser-Based Bioprinting
Stereolithography (SLA) is an AM technique that uses ultraviolet (UV) or visible light to
cure photosensitive polymers in a layer-by-layer fashion, as shown in Figure 3. This nozzle-free
technique eliminates the negative effects of shear pressure encountered when using nozzle-based
bioprinting. It offers a fast and accurate fabrication, with resolutions ranging between 5–300
µ
m [
61
,
62
].
Polymerization occurs at the top of the bioink vat where the biomaterial is exposed to the light energy.
After each layer is polymerized, the platform supporting the structure will be lowered in the vat,
enabling a new layer to be photopolymerized on top.

Materials 2018,11, 2199 9 of 21
Materials 2018, 11, x FOR PEER REVIEW 9 of 22
Figure 3. Schematic of Stereolithography Bioprinting. Photopolymerization occurs on the surface of
the vat where the light-sensitive bioink is exposed to light energy. Axial platform moves downward
the Z-axis during fabrication. This layer-by-layer technique does not depend on the complexity of the
design, rather on its height.
Photoinitiators are chemical molecules that create reactive agents when exposed to light energy,
which react with monomers of a material to then initiate the formation of polymer chains.
Photoinitiators are sensitive to different ranges of wavelength; some are triggered by UV and others
by visible light. The stiffness and network density of the cured resin depends on the concentration of
the photointiator but higher concentrations might exhibit adverse cytotoxic effects. However,
different photoinitiators have different cytotoxicity levels. The most commonly used and the least
cytotoxic photoinitiators are Irgacure 2959 for UV cross-linkage and eosin Y for visible light [63].
Eosin Y has even shown to be less toxic than Irgacure 2959 [63]. UV light will affect cells and
introduce mutations [64]; therefore, visible light-based photocross-linkage has been adopted more
frequently in SLA as well as in situ applications [65,66]. Photopolymerization is also employed
during or post-fabrication via inkjet- and extrusion-based printing to harden the prints [26,57].
Due to the risk of damaging the cells through the use of UV light or cytotoxic effects of the
photoinitiators, several researchers have investigated alternative means to enable
photopolymerization of bioinks. Hoffmann et al. developed a class of materials that crosslink
without the presence of a photoinitiator using a thiol-ene reaction [67]. The used monomers
comprise two classes of monomers containing at least two alkene or thiol groups. These two
components react spontaneously under ultraviolet (UV)-irradiation at a wavelength of
approximately 266 nm. A 1:1 ratio of thiol and alkene exhibited high cell viability after 3 days, ≈95%.
However, doubling the thiol content resulted in a cytotoxic effect, even though this amount of thiol
groups provides high amounts of surface functional groups, allowing greater subsequent surface
functionalization.
Zhang et al. used UV laser in the form of Bessel beam [68]. Bessel beam does not diffract and
spread out, which will be useful to increase print fidelity and decrease fabrication time. The
precursor hydrogel was prepared from GelMA, PEGMA and Irgacure 2959. Human umbilical vein
endothelial cells (HUVECs) were encapsulated in the hydrogel. Cell-laden fibers with diameters 25,
43 and 75 µm were fabricated and cell viability was 95% after 3 days. This technique has potential in
Figure 3.
Schematic of Stereolithography Bioprinting. Photopolymerization occurs on the surface of
the vat where the light-sensitive bioink is exposed to light energy. Axial platform moves downward
the Z-axis during fabrication. This layer-by-layer technique does not depend on the complexity of the
design, rather on its height.
Photoinitiators are chemical molecules that create reactive agents when exposed to light energy,
which react with monomers of a material to then initiate the formation of polymer chains. Photoinitiators are
sensitive to different ranges of wavelength; some are triggered by UV and others by visible light. The stiffness
and network density of the cured resin depends on the concentration of the photointiator but higher
concentrations might exhibit adverse cytotoxic effects. However, different photoinitiators have different
cytotoxicity levels. The most commonly used and the least cytotoxic photoinitiators are Irgacure 2959 for UV
cross-linkage and eosin Y for visible light [
63
]. Eosin Y has even shown to be less toxic than Irgacure 2959 [
63
].
UV light will affect cells and introduce mutations [
64
]; therefore, visible light-based photocross-linkage has
been adopted more frequently in SLA as well as in situ applications [
65
,
66
]. Photopolymerization is also
employed during or post-fabrication via inkjet- and extrusion-based printing to harden the prints [26,57].
Due to the risk of damaging the cells through the use of UV light or cytotoxic effects of the
photoinitiators, several researchers have investigated alternative means to enable photopolymerization
of bioinks. Hoffmann et al. developed a class of materials that crosslink without the presence of
a photoinitiator using a thiol-ene reaction [
67
]. The used monomers comprise two classes of monomers
containing at least two alkene or thiol groups. These two components react spontaneously under
ultraviolet (UV)-irradiation at a wavelength of approximately 266 nm. A 1:1 ratio of thiol and alkene
exhibited high cell viability after 3 days,
≈
95%. However, doubling the thiol content resulted in
a cytotoxic effect, even though this amount of thiol groups provides high amounts of surface functional
groups, allowing greater subsequent surface functionalization.
Zhang et al. used UV laser in the form of Bessel beam [
68
]. Bessel beam does not diffract and
spread out, which will be useful to increase print fidelity and decrease fabrication time. The precursor
hydrogel was prepared from GelMA, PEGMA and Irgacure 2959. Human umbilical vein endothelial
cells (HUVECs) were encapsulated in the hydrogel. Cell-laden fibers with diameters 25, 43 and 75
µ
m
were fabricated and cell viability was 95% after 3 days. This technique has potential in fabricating
tubular constructs and porous scaffolds under a shortened fabrication time; however, is limited to low
structural complexity.
Tuan et al. developed a visible light-based stereolithography using Lithium
phenyl-2,4,6-trimethylbenzoylphosphinate (LAP) [
69
], which is a UV-sensitive photoinitiator
that can also respond to near-UV blue light [
63
]. Human adipose-derived stem cells (hADSCs) were
suspended in a Poly(ethylene glycol) diacrylate (PEGDA)/LAP solution. Although near-UV blue light,

Materials 2018,11, 2199 10 of 21
400–490 nm can be damaging to mammalian cells [70], after fabrication the hADSCs exhibited a high
metabolic activity, increasing by 75% and 50% after 5 and 7 days, respectively.
Other photoinitiators that can absorb visible light are camphorquinone and eosin Y, that crosslink
at wavelengths of 400–700 nm and 514 nm, respectively [
63
]. Wang et al. mixed PEG with eosin Y and
methacrylated gelatin (GelMA). Samples without GelMA exhibited decreased cell viability compared
to the samples consisting of 5% and 7.5% GelMA, which maintained cell viabilities of ~80% after
5 days [
71
]. The slightly decreased cell viability could be related to the fact that PEG is non-adhesive,
causing the death of anchorage-dependent cells [72,73].
Wang et al. fabricated GelMA-based scaffolds via visible light-based SLA [
74
]. The precursor gel
was mixed with eosin Y and NIH-3T3 fibroblasts. The scaffolds were crosslinked by a commercial
projector at 522 nm wavelength. After 5 days in culture, most of the cells adhered to bioink.
Hu et al. studied the cytotoxicity of chitosan-based scaffolds that were mixed with either
camphorquinone, fluorescein or riboflavin [
75
]. Fluorescein and riboflavin are blue light-absorbing
initiators. Camphorquinone exhibited relatively low cell viability, ~40%, whilst the other two
photoinitiators exhibited cell viabilities >80%. Camphorquinone is more commonly used than the other
two photoinitiators; however, biocompatibility results of camphorquinone have been inconsistent in
literature [75–77].
Stereolithography has much to offer in its application to bioprinting. The absence of shear
stress and no limitation on bioink viscosity make it as an appealing choice for incorporating cells
within scaffolds. However, the limitations of SLA include the damage caused by UV and near UV
light to cell DNA, the limited choice of photosensitive biomaterials as well as the cytotoxicity of
added. Some researchers have already begun to look for alternatives, such as using photoinitiator-free
materials or visible light-absorbing photoinitiators [67,78].
4. Laser-Assisted Bioprinting
Laser-assisted printing was initially developed to deposit metals onto receiver sheets [
79
,
80
].
Odde and Renn later developed the technique to print viable embryonic chick spinal cord cells [
81
].
Laser-assisted bioprinting (LAB) consists of three parts: a donor-slide (or ribbon), a laser pulse and
a receiver-slide. A ribbon is made of a layer of transparent glass, a thin layer of metal, and a layer of
bioink. The bioink is transferred from the ribbon onto the receiver slide when the metal layer under
the hydrogel is vaporized by a laser pulse, as depicted in Figure 4. This scaffold-free technique has
very high cell viabilities (>95 [
54
]) and a resolution between 10–50
µ
m [
1
]. Some studies using LAB
have demonstrated an accuracy of a singular cell per droplet [82].
Materials 2018, 11, x FOR PEER REVIEW 11 of 22
Figure 4. Schematic of Laser-assisted Bioprinting. (a) transparent glass, (b) thin metal layer,
(c) vaporization-induced bubble. Bubble nucleation induced by laser energy propels droplets of
bioink towards the substrate. This technique has minimal effect on cell viability. A receiver-slide can
be a biopaper, polymer sheet or scaffold.
Gruene et al. conducted a study to observe the effects of the LAB laser pulse had on printed
mesenchymal stems cells (MSCs). It was found that the laser pulse had a negligible effect. There
were no reported changes in gene expression caused by the heat shock of the laser pulse, and cell
proliferation rates were as high as the control of non-printed cells after 5 days in cell culture [24].
Alkaline phosphatase (ALP) expression and calcium accumulation were similar to non-printed
MSCs after 3 weeks in osteogenic medium.
Keriquel et al. printed in situ MSCs on to a collagen/nanohydroxyapatite (nHA) disks placed
cranial defects [82]. Compared to acellular collagen/nHA disks, the disks with the bioprinted MSC
cells exhibited a larger bone volume after 2 months. Michael et al. printed 20 layers of keratinocytes
on top of 20 layers of fibroblasts, situated on top of a carrier matrix, Matriderm
®
that provided
stability [21]. Keratinocytes developed into a stratified dense tissue in an in vivo study after 11 days
implanted subcutaneously in mice, and demonstrated the potential for LAB in skin tissue
regeneration.
LAB has the ability to position multiple cell types with a high degree of accuracy, with several
studies demonstrating singular the capability of positioning a singular cell per droplet [29,81,83].
However, it is an expensive process to perform and suffers from low stability and scalability. It has
shown great potential when combined with other biofabrication techniques [29,84].
5. Extrusion-Based Bioprinting
Extrusion-based printing is a pressure-driven technology. The bioink is extruded through a
nozzle, driven either by pneumatic or mechanical pressure, and deposited in a predesigned
structure, as depicted in Figure 5 [50]. The main advantage of extrusion bioprinting is the ability to
print with very high cell densities [85,86]. Despite its versatility and benefits, it has some
disadvantages when compared to other technologies. The resolution is very limited, as a minimum
feature size is generally over 100 µm, which is a poorer resolution than that of other bioprinting
techniques [87]. This could limit its application for certain soft tissue applications that require small
pore sizes for an improved tissue response [11,86,88], however could still be applicable to hard
tissues with size larger than 10 mm [35,86]. The pressure used for the extrusion of the material has
Figure 4.
Schematic of Laser-assisted Bioprinting. (
a
) transparent glass, (
b
) thin metal layer,
(
c
) vaporization-induced bubble. Bubble nucleation induced by laser energy propels droplets of
bioink towards the substrate. This technique has minimal effect on cell viability. A receiver-slide can be
a biopaper, polymer sheet or scaffold.

Materials 2018,11, 2199 11 of 21
Gruene et al. conducted a study to observe the effects of the LAB laser pulse had on printed
mesenchymal stems cells (MSCs). It was found that the laser pulse had a negligible effect. There were
no reported changes in gene expression caused by the heat shock of the laser pulse, and cell
proliferation rates were as high as the control of non-printed cells after 5 days in cell culture [
24
].
Alkaline phosphatase (ALP) expression and calcium accumulation were similar to non-printed MSCs
after 3 weeks in osteogenic medium.
Keriquel et al. printed in situ MSCs on to a collagen/nanohydroxyapatite (nHA) disks placed
cranial defects [
82
]. Compared to acellular collagen/nHA disks, the disks with the bioprinted MSC
cells exhibited a larger bone volume after 2 months. Michael et al. printed 20 layers of keratinocytes
on top of 20 layers of fibroblasts, situated on top of a carrier matrix, Matriderm
®
that provided
stability [
21
]. Keratinocytes developed into a stratified dense tissue in an
in vivo
study after 11 days
implanted subcutaneously in mice, and demonstrated the potential for LAB in skin tissue regeneration.
LAB has the ability to position multiple cell types with a high degree of accuracy, with several
studies demonstrating singular the capability of positioning a singular cell per droplet [
29
,
81
,
83
].
However, it is an expensive process to perform and suffers from low stability and scalability. It has
shown great potential when combined with other biofabrication techniques [29,84].
5. Extrusion-Based Bioprinting
Extrusion-based printing is a pressure-driven technology. The bioink is extruded through
a nozzle, driven either by pneumatic or mechanical pressure, and deposited in a predesigned structure,
as depicted in Figure 5[
50
]. The main advantage of extrusion bioprinting is the ability to print with
very high cell densities [
85
,
86
]. Despite its versatility and benefits, it has some disadvantages when
compared to other technologies. The resolution is very limited, as a minimum feature size is generally
over 100
µ
m, which is a poorer resolution than that of other bioprinting techniques [
87
]. This could
limit its application for certain soft tissue applications that require small pore sizes for an improved
tissue response [
11
,
86
,
88
], however could still be applicable to hard tissues with size larger than
10 mm [
35
,
86
]. The pressure used for the extrusion of the material has the potential to alter the cell
morphology and function, although several studies have reported [
86
]. Overall, before printing of
the hydrogel can be performed a detailed study with different process parameters including viscosity,
nozzle diameter and the accompanied shear stress has to be evaluated [
89
,
90
]. This fabrication
technique uses highly viscous hydrogel and does not necessarily require any chemical additives
for the curing of printed structure [
86
]. Rheological behavior of the hydrogel ink is very important
for extrusion-based bioprinting. Hydrogels are mostly non-Newtonian fluids, meaning that their
viscosity changes with shear rate. However, the more viscous the bioink, the higher the induced
shear-stress during printing, resulting in higher cell apoptotic activity. An important phenomenon
in non-Newtonian fluids is shear thinning, which is a drop of viscosity with an applied shear force.
This has a direct impact on the print quality, enabling a plug-like flow to be established, providing
greater control over starting and stopping the extrusion process [
91
]. Although low viscosities result
in less dense networks that could allow for better cellular infiltration, too low viscosities will produce
a structure that has a poor definition that will ultimately affect print fidelity.

Materials 2018,11, 2199 12 of 21
Materials 2018, 11, x FOR PEER REVIEW 12 of 22
the potential to alter the cell morphology and function, although several studies have reported [86].
Overall, before printing of the hydrogel can be performed a detailed study with different process
parameters including viscosity, nozzle diameter and the accompanied shear stress has to be
evaluated [89,90]. This fabrication technique uses highly viscous hydrogel and does not necessarily
require any chemical additives for the curing of printed structure [86]. Rheological behavior of the
hydrogel ink is very important for extrusion-based bioprinting. Hydrogels are mostly
non-Newtonian fluids, meaning that their viscosity changes with shear rate. However, the more
viscous the bioink, the higher the induced shear-stress during printing, resulting in higher cell
apoptotic activity. An important phenomenon in non-Newtonian fluids is shear thinning, which is a
drop of viscosity with an applied shear force. This has a direct impact on the print quality, enabling a
plug-like flow to be established, providing greater control over starting and stopping the extrusion
process [91]. Although low viscosities result in less dense networks that could allow for better
cellular infiltration, too low viscosities will produce a structure that has a poor definition that will
ultimately affect print fidelity.
Figure 5. Schematic of Extrusion-based Bioprinting; from left, pneumatic-based and right,
mechanical-based. Struts are extruded via pneumatic or mechanical pressure through micro-nozzles.
Extrusion-based techniques can produce structures with great mechanical properties and print
fidelity.
A study conducted by Chung et al., observed the bioink properties and the printability of
alginate-gelatin blends. Using Alg-Gel ink solutions, the printing of scaffolds from three different
alginate concentrations (1, 2, and 4% w/v) were compared. Both printed scaffolds using 2% Alg-Gel
and 4% Alg-Gel demonstrated defined structures and maintained their ability to support optimal
cell growth. The highly hydrated network structure permits the exchange of gases and nutrients
[92]. When choosing a hydrogel to use as the base material, a trade-off must be made between
rigidness and softness in order to have a strong supporting structure that allows for nutrient
infiltration and the capability to encapsulate cells. High concentrations or crosslink densities are
needed to keep a good printing fidelity, yet this limits cell migration. However, low concentrations
usually have a poor printability and low mechanical properties. To improve the mechanical
properties of the hydrogel, reinforcing fibers like PCL can be used [93]. Photopolymerization is
emerging as a promising crosslinking reaction for bioprinting because it enables the rapid formation
Figure 5.
Schematic of Extrusion-based Bioprinting; from left, pneumatic-based and right,
mechanical-based. Struts are extruded via pneumatic or mechanical pressure through micro-nozzles.
Extrusion-based techniques can produce structures with great mechanical properties and print fidelity.
A study conducted by Chung et al., observed the bioink properties and the printability of
alginate-gelatin blends. Using Alg-Gel ink solutions, the printing of scaffolds from three different
alginate concentrations (1, 2, and 4% w/v) were compared. Both printed scaffolds using 2% Alg-Gel
and 4% Alg-Gel demonstrated defined structures and maintained their ability to support optimal
cell growth. The highly hydrated network structure permits the exchange of gases and nutrients [
92
].
When choosing a hydrogel to use as the base material, a trade-off must be made between rigidness
and softness in order to have a strong supporting structure that allows for nutrient infiltration
and the capability to encapsulate cells. High concentrations or crosslink densities are needed to
keep a good printing fidelity, yet this limits cell migration. However, low concentrations usually
have a poor printability and low mechanical properties. To improve the mechanical properties
of the hydrogel, reinforcing fibers like PCL can be used [
93
]. Photopolymerization is emerging
as a promising crosslinking reaction for bioprinting because it enables the rapid formation of
hydrogels immediately after printing to maintain print fidelity through the incidence of light energy
at appropriate wavelengths [
1
,
94
–
97
]. The printing resolution can also affected by the diffusion and
fusion of the bioinks, which could be solved by reducing the extrusion rate or accelerating the moving
speed. With good cell compatibility of the hydrogel material and the high printing quality with
appropriate printing process parameters, the hydrogel deposition in the fabrication of tissues or organs
can be obtained [98].
An important characteristic for the hydrogel is that it should maintain its mechanical properties
after printing. During printing, the hydrogel is subjected to different forces. In nozzle based printing
systems, such as with inkjet and extrusion-based techniques, high shear forces can break or disrupt
the interlinking bonds of the hydrogel molecular network. This damage to the hydrogel crosslinking
can cause a drop in viscosity and a reduction in print fidelity. To overcome this issue, research has
been conducted into self-healing hydrogels [
99
]. A self-healing hydrogel can retain its printed shape
due to its non-covalent reversible bonds [
100
,
101
]. An improved structure of hydrogels is a structure
that has interpenetrating polymer networks (IPNs, which consist of 2 (or more) polymer networks;
where one is crosslinked in the immediate presence of another [
102
]. The networks can be crosslinked
simultaneously or sequentially, from heterogeneous or homogeneous materials. An example of
IPNs made of heterogeneous materials is double network (DN) IPNs, which is fabricated in a 2-step

Materials 2018,11, 2199 13 of 21
polymerization process of rigid and soft hydrogels [
103
]. Biocompatible DNs have been successfully
employed in cell encapsulation [104].
Cell survivability and function can also be negatively influenced by the extrusion process.
In highly concentrated bioinks, shear stresses have the potential to cause cell apoptosis and a drop in
the number of living cells [
1
,
86
,
105
]. Shear stress can also affect cell morphology and metabolic activity,
as well as the adhesiveness of the cells to the substrate [
86
]. However, the overall cellular response is
dependent upon cell type, as some cells are more resistant than others [86].
Extrusion printing can be regarded as a promising technology that allows the fabrication of
organized constructs at clinically relevant sizes within a reasonable time frame. However, selection of
biomaterial and bioink concentration is important for the survival of the cells during fabrication,
as well as the maintenance of cell viability and functionality post-printing.
Lee et al. used an extrusion bioprinter to regenerate an ear formed of auricular cartilage and fat
tissue [
106
]. The ear shaped scaffold was fabricated using chondrocytes and adipose-derived stromal
cells, encapsulated in a hydrogel composed of PCL and poly(ethylene-glycol) (PEG). The bioprinted
ear achieved a 95% cell viability [
106
]. The regeneration of the ear has been considered to be
a challenge due to its complex structure and composition, which is difficult to replicate using traditional
fabrication techniques.
Kundu et al. produced cartilage scaffolds by extruding alginate hydrogel onto PCL [
107
].
Scaffold were printed either with or without human inferior turbinate-tissue derived mesenchymal
stromal cells (hTMSCs) within the alginate bioink. Better chondrogenic function was observed when
the hTMSCs were encapsulated in alginate gel as well as an increase in extra cellular matrix (ECM)
production without an adverse tissue response when implanted into the dorsal subcutaneous spaces
of mice [
107
]. The encapsulation of the cells in alginate hydrogel showed negligible effects on the
viability of the chondrocytes which addressed the formation and synthesis of cartilaginous ECM.
Pati et al. developed a hybrid scaffold combining PCL and decellularized extracellular matrix
(dECM) [
108
]. The dECM bioink was loaded with stem cells derived from adipose, cartilage and
heart tissues, and deposited into a PCL framework. It was observed that there was a cell-to-cell
interconnectivity within 24 h and a cell and viability of 90% on day 7. This study shows the ability
to print complex structures with appropriate material and cells, which can provide an optimized
microenvironment that is conductive to the growth of 3D structured tissues.
Miri et al. demonstrated the possibility to create hierarchical cell laden structures to mimic
multicellular tissues [
26
]. For
in vitro
studies, hydrogels including poly(ethylene glycol) diacrylate
(PEGDA) and methacrylated gelatin (GelMA) loaded with NIH/3T3 fibroblasts and C2C12 skeletal
muscle cells were printed into structures resembling musculoskeletal junctions, muscle strips and
tumor angiogenesis. The prints retained interfaces and adequate proliferation rates after 3, 5 and
7 days in cell culture. PEGDA-framed chips that had a concentration-gradient of GelMA ranging
from 5–15%, were implanted subcutaneously in rats. The result showed formation of the blood vessel
network in the bioactive GelMA hydrogels, while the PEGDA served as the frame in the bioprinted
multimaterial structure. This novel pneumatic-based process of creating microfluidic devices enabled
the printing of different cell suspensions in order to achievemultimaterial devices.
Extrusion bioprinting is a promising technique to create biomimetic structures to replace tissues
and organs. This technique was also efficient in creating microfluidic chips for research applications.
Despite its great versatility and feasibility in vertical printing, extrusion-based bioprinting has
a relatively limited resolution that does not allow for cell positioning, and requires an advanced
hydrogel bioink that maintains cell viability as well as mechanical integrity which has led to the
development and use of self-healing hydrogels as well as interpenetrating polymer networks.
6. Discussion
3D bioprinting is a relatively new aspect to tissue engineering and has opened the possibility
of creating an unprecedented biomimicry, which could ultimately replace the current gold standard

Materials 2018,11, 2199 14 of 21
of autografts. Biomimicry, in form and function, has great significance in regenerative medicine,
drug screening and understanding pathology [
109
].
In vitro
applications have been used to assess
pathological and toxicological conditions, as well as implant integration, and offers a methodology
with a high-throughput [
110
]. Biomimetic microfluidic chips have great potential in replacing animal
studies for drug and material screening.
Each bioprinting technique has different requirements for the bioink that can create diverse effects
on the encapsulated cells. Inkjet bioprinting provides high resolution and accurate cell positioning.
However, it requires the bioink to have a low concentration, which may result in poor structural
integrity and inefficient cell encapsulation. This technique has shown great success in creating neural
and skin tissues [
32
,
59
]. In skin tissue engineering, scaffolds fabricated using inkjet bioprinting have
delivered better results when compared to a commercial graft Alpigraf
®
to repair full thickness wounds
in mice [37].
Stereolithography offers the possibility of printing cell-laden structures with the shortest
fabrication time possible, hence limiting the exposure of the cells to non-physiological conditions.
SLA fabrication does not inflict shear stresses upon the cells, unlike in nozzle based techniques, which
have the potential to cause cell apoptosis. However, complex designs that include hollow structures
(vessels, vasculature or ducts), can become blocked due to remnants of the precursor hydrogel within
the printed pores [
26
]. Another problem with SLA is that surplus bioink is used as fabrication is
performed in a vat. That vat is filled with a larger volume of biomaterial, cells and biomolecules than
what is needed for the fabrication of the scaffold.
Extrusion-based printing is the most feasible technique in terms of vertical configuration, although
has the lowest reported cell survival among all techniques. The low survivability is due to the shear
stress that arises during printing. An important aspect of extrusion printing is its influence on the
hydrogel during and after printing. Due to the high shear stresses induced during printing it is
possible that the hydrogel could lose its structural integrity. This has led to the development of
self-healing hydrogels, which regain their mechanical integrity after the application of shear [
111
].
Extrusion-based bioprinting has succeeded in creating complex tissue constructs and multi-material
microfluidic devices [36,39].
A problem encountered by all techniques when using photopolymerization to harden the bioink, is
the cytotoxicity of the photoinitiators used and the damage inflicted by UV (10–400 nm) or near-UV blue
(400–490 nm) irradiation. However, alternatives to the use of UV light and the use of photoinitiators
are under investigation. Visible light-sensitive photoinitiators have reported less cytotoxicity than the
most commonly used UV-sensitive photoinitiators [63], as well as an enhanced print fidelity [78].
Post-fabrication, cell-laden scaffolds can be incubated in culture medium to ensure the attachment
of cells [
112
]. Incubation for longer periods (21 days) has resulted in an increase of mechanical strength
of the scaffolds due to tissue development [
57
]. Incubation can be static in cell culture or dynamic
using bioreactors. Dynamic culturing can provide continuous infiltrating flow of medium and/or
compressive/tensile loading, which is most beneficial for cartilage and bone tissue engineering [
113
].
Current research demonstrates the feasibility and efficiency of using more than one fabrication
technique in the manufacturing process. Inkjet printing and LAB have the capability of accurate
cell positioning with both of them having achieved the positioning of singular cells per droplet.
However, inkjet printing is limited by its ability to produce a 3D architecture, whereas LAB only
positions the bioink onto a prefabricated scaffold and is also associated with a high cost. In contrast,
extrusion bioprinting has fast fabrication times for large 3D structures, yet has poor cell survivability.
Therefore, by combining either inkjet bioprinting or LAB with extrusion printing could provide the
ideal combination for producing a scaffold that has both physiologically relevant proportions as well
as supports viable cells.
Research has already been implemented combining different printing techniques. In a study
by Kim et al., a skin model was fabricated using an extrusion printer to create the main supporting
structure and an inkjet printer was used to position dermal fibroblasts and epidermal keratinocytes

Materials 2018,11, 2199 15 of 21
within the scaffold [
114
]. The bioprinted scaffold formed dermal and epidermal layers after culturing.
Another study combined extrusion printing with stereolithography to create a model for cancer
research, where microfluidic devices were fabricated using a digital micro-mirror device and pneumatic
extrusion, to understand tumor angiogenesis [
26
]. In situ applications, where the cell-laden biomaterial
is directly deposited into the defect, are also being investigated for accelerated wound healing and
bone regeneration, which have demonstrated improved results in comparison to non-cell containing
grafts [53,65,66].
Finally, another aspect of bioprinting is its potential to provide prevascularization of the scaffolds.
Accurate cell positioning in LAB and inkjet bioprinting techniques could enable a vasculature to be
printed into a scaffold. Both techniques have shown promising results in positioning endothelial
cells to induce angiogenesis [
29
,
42
]. Prevascularization is essential to avoid necrotic failure of the
implantation. Other cell positioning research based on inkjet techniques shows great potential in
constructing neural networks within large structures [32].
7. Conclusions
Additive manufacturing has been heavily applied to tissue engineering over the past decade.
Bioprinting enables the production of scaffolds with a homogeneous distribution of cells throughout
a scaffold. An organized distribution of different cell types can be positioned within the supporting
material, mimicking tissues with multiple cell types or the interface between two tissues. While the
choice of material and design impact the viability and proliferation of the printed cells, the different
techniques have also shown variable cell activities post-fabrication. Bioprinting is still under
development and has many bridges to cross before entering the clinical world, particularly as an
in situ direct application. From this brief review, it is concluded that different applications require
different fabrication techniques, depending on required resolution, speed, cost, the ability to print
vertically etc. Future developments are now concentrating on the combining of techniques to work in
a complementary fashion to optimize the process of creating tissue-mimicking structures.
Author Contributions:
S.A., S.R. and Z.I. conducted a literature review to provide the information of for this
review article, Z.P.K., P.M.R. and M.B. wrote the article, R.S. and O.J. proof read the manuscript and helped with
the final editing.
Funding: This research received no external funding
Conflicts of Interest: The authors declare no conflict of interest.
Abbreviations
AFS Amniotic fluid-derived stem cells
AG Agarose
ALP Alkaline phosphatase
AM Additive manufacturing
ATCC Mouse neural stem cell lines
BMSCs Bone marrow stromal cells
BrCa Breast cancer cells
CAD Computer aided design
CT Computer Tomography
dECM Decellularized extracellular matrix
DN Double network
DNA Deoxyribonucleic acid
ECM Extracellular matrix
GelMA Gelatin methacryloyl
HA Hydroxyapatite
hADSCs Human adipose-derived stem cells
HAMa Hyaluronic acid–methacrylate
HMECs Human microvascular endothelial cells

Materials 2018,11, 2199 16 of 21
HMVECs Human dermal microvascular endothelial cells
Hs68 Human dermal fibroblasts
hTMSCs Human inferior turbinate-tissue derived mesenchymal stromal cells
HUVECs Human umbilical vein endothelial cells
IPFP Human infrapatellar fat pad derived adipose stem cells
IPNs Interpenetrating polymer networks
LAB Laser-assisted bioprinting
LAP Lithium phenyl-2,4,6-trimethylbenzoylphosphinate
MRI Magnetic Resonance Imaging
MSCs Human bone marrow mesenchymal stem cells
nHA Nanocrystalline hydroxyapatite
NHDFs Human dermal fibroblasts
NHEKs Neonatal human epidermal keratinocytes
PCL Polycaprolactone
PEG Poly(ethylene-glycol)
PEGDA poly(ethylene glycol) diacrylate
PEGDMA Poly(ethylene glycol) dimethacrylate
PLA Polylactide fibers
PVA polyvinyl alcohol
SA Sodium alginate
SLA Stereolithography
UV Ultraviolet
VEGF Vascular endothelial growth factor
β-TCP Beta-tricalcium phosphate
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