The recent development and applications of fluidic channels by 3D printing
Zhou Journal of Biomedical Science
The recent development and applications of fluidic channels by 3D printing
0 Singapore Centre for 3D Printing (SC3DP), School of Mechanical and Aerospace Engineering, Nanyang Technological University , 50 Nanyang Ave , Singapore 639798 , Singapore
The technology of “Lab-on-a-Chip” allows the synthesis and analysis of chemicals and biological substance within a portable or handheld device. The 3D printed structures enable precise control of various geometries. The combination of these two technologies in recent years makes a significant progress. The current approaches of 3D printing, such as stereolithography, polyjet, and fused deposition modeling, are introduced. Their manufacture specifications, such as surface roughness, resolution, replication fidelity, cost, and fabrication time, are compared with each other. Finally, novel application of 3D printed channel in biology are reviewed, including pathogenic bacteria detection using magnetic nanoparticle clusters in a helical microchannel, cell stimulation by 3D chemical gradients, perfused functional vascular channels, 3D tissue construct, organ-on-a-chip, and miniaturized fluidic “reactionware” devices for chemical syntheses. Overall, the 3D printed fluidic chip is becoming a powerful tool in the both medical and chemical industries.
Fluidic channel; Lab-on-a-chip; 3D printing; Diagnosis; Tissue engineering; Reactionware
Shrinking the bulky and costly laboratory equipment into a
single small, user-friendly, easily replicable chip provides a
significant advantage over traditional assays. Lab-on-a-chip
(LOC) technologies or micro-total analysis systems (μTAS)
with microfluidics have been continuously evolving from
simple single-function devices to analytical systems with
multiple functionalities and revolutionizing the research
fields of chemistry, physics, pharmacology, cell biology,
chemical biology, neuroscience, biomechanics, bioanalysis,
and tissue engineering [
]. Particularly, they have
ubiquitous presences in various clinical and forensic analysis [
such as cell sorting and isolation [
], cellular analysis ,
biosensor and point-of-care (POC) diagnosis [
pharmacological screening [
2, 11, 12
], proteomics and metabolomics
, immunoassays [
], genetic analysis or genomics [
multi-cellular tissue spheroid fabrication, organ-on-a-chip
using different tissue spheroids [
], and bioreactor for
co-culture and maturation of micro-organs. Current
microfluidic systems have extended their applications with the
integration of several functions, for example, cell/tissue
incubation, enzymatic processing, biochemical analysis,
optoelectronic measurement, and computer-controlled
microfluidics. In comparison to the traditional macro-scale
methods, these microfluidic chips have the capabilities of (1)
streamlining complex assay protocols, (2) minimizing the
sample and reagent volumes, (3) maximizing the
measurement of precious sample at the reduce performance time,
power consumption, and substantial cost, (4) accurately
manipulating the cell microenvironment, and (5) providing
scalability and batch screening of multiple samples in a
massive parallel style. Despite the rapid development of
microfluidics and applications in biological research and
biomedical engineering over the past decades, the wide and
practical acceptance (e.g., commercial point-of-care testing)
is still not very satisfactory. One of the reasons may be the
requirement of a highly adaptable, rapid, and easy process of
fabricating the microfluidic systems with increasing
prevalence and complexity and the absence of a “killer application”
that would outperform existing traditional methods.
3D microfluidic chips can overcome the limitations of
conventional 2D designs and have the potential
advantages of improved observation efficiency [
3D motion [
], and integration of more functions,
especially for patterning with liquids. With the increases in the
complexity and more sophisticated tasks, the transport of
different fluid streams becomes easy in 3D configuration.
However, the complexity of manufacturing such systems
(e.g., more steps for pattern fabrication, alignment, and
sealing in photolithography in comparison to the 2D
configuration) has deterred their wide use. There are several
approaches for mass replication and production of 3D
microfluidic devices, for example, micro-machining,
casting, hot embossing, in situ construction injection, and
laser ablation [
]. Some of them require large equipment
space, intensive labor, time consumption and suffer from
limited bio-compatible materials. It is troublesome to
produce multiple photomasks in high resolution (<10 μm)
and align and expose sequential layers of photoresist for
the soft lithography. The need for fabricating the
replication master limits its application on a smaller scale, for
example, academic research and biological applications. In
addition, the equipment complexity and operator’s skills
are relatively high. Therefore, a quick, easy, and direct
fabrication is preferred for the end users. Recently,
advances in 3D printing may simplify the fabrication
process of fluidic devices into a single step [
Additive manufacturing or 3D printing could produce
complicated intricate architectures effectively at
relatively low cost and infrastructure investment, but high
attainability. Freedom of product design for
manufacturing and assembly greatly encourages product innovation
universally available to any customers by creating unique
bespoke one-off objects. This technology has witnessed
an explosive growth in the manufacturing industry and
the consumer market. Because no etching or dissolution
is required, the process of adding materials is
environmentally friendly and economically efficient. The global
market for 3D printing has increased to $4.1 billion in
2014 and may reach $20.2 billion by 2020 [
3D printed structures cannot compete with those
manufactured by photolithography in the resolution now, they
enable the enhanced geometrical control of the
unprecedented channel shape and complexity (e.g., cross-section
and height) quickly and inexpensively that has been
previously impossible. In comparison to the conventional
clean-room-based photolithography, it offers several
major advantages: (1) rapid prototyping and replication
of products with its attendant benefits and full
automation to positively disrupt the development cycles, (2)
greatly simplifying the manufacturing process without a
replication master, assembly, and extensive labor, (3) no
requirement of clean-room environment for
comparatively low cost in manufacturing infrastructure, (4)
arbitrary channel shapes instead of rectangular one in
photolithography, (5) very simple procedures for
structures with various heights in a single step instead of
using the layer-by-layer strategy, (6) molding suspended
structures without any alignment or sacrificial parts, (7)
multiple materials for various applications (e.g., artificial
tissue scaffolds), (8) dramatically lowering the barrier to
creating sophisticated 3D biomedical models, (9) the
great topography flexibility of multiple 2D layers stacked
together, and (10) enabling users to adopt a “fail fast and
often” strategy [
In this paper, the recent advances in 3D printing in the
fluidic channel and its applications in biology and
biomedical engineering are reviewed. First, all materials used in
the fabrication are summarized. Then, various
manufacturing approaches, such as stereolithography (SL), two
photon polymerization (2PP), fused deposition modeling
(FDM), polyjet, and 3D bioprinting, are introduced. The
strategy of removing the scaffold made by 3D printing also
enables the fabrication of channels in high complexities
and throughput. Currently investigated applications, such
as the deposition and detection of cells and proteins,
development of bacterial communities, formation of the
microvascular network, stimulation of cell growth, and
construction of 3D tissue/organ, are listed. Finally, the
trend of 3D printing and fluidic channel in the
lab-on-achip is discussed, and the improvement on the current
limitations is necessary for the fast commercialization and
wide acceptance of “killer-applications”.
Materials used for the fabrication of a microfluidic
system require the consideration of their function, the
degree of integration, and biological application, for
example, cellular compatibility, supportability (e.g.,
oxygen and nutrient diffusion), optical transparency, and
mechanical properties. The most popular materials in
molding approaches are polydimethylsiloxane (PDMS)
and thermoplastics. Devices molded in thermoplastics
[e.g., polystyrene, polymethyl methacrylate (PMMA),
polyurethane] enable higher throughput but do not
necessarily allow superior manufacturability. However,
thermoplastics and injection molding are not amenable
to rapid-prototyping because both the equipment and
the molds are expensive, the turn-around times for the
fabrication of metallic molds can be on the order of
weeks, and the molding procedure requires substantial
expertise. Plastics do not have the high gas solubility as
PDMS which obeys Henry’s law.
PDMS is usually selected because of its (1) gas
permeability for keeping cells and bacteria alive for a long time,
(2) elasticity (Young’s modulus of 2 MPa) [
], capable of
making micro-pumps and valves approximately 1000
times smaller than that of hard plastics, (3) simple
chemical modification using well-known silane chemistry, (4)
optical transparency at the wavelength of 240–1100 nm
], nontoxic, electrically insulating, and impermeability
to liquids, (5) conformal and easy to mold with high
fidelity and precision (in the order of 10 nm) [
], (6) fairly
low cost, (7) free copyright, (8) biocompatible, and (9)
rapid prototype using simple procedures. However, with
an increasing focus on translation and low-cost devices,
molding approaches illustrate their shortcomings: (1)
PDMS molding (curing, assembly, bonding, and inlet
punching) using photolithography is substantially
laborintensive and complex for inexperienced people [
that it is hard to fully automate and disseminate out of
research labs for commercialization or large-scale
]; (2) the user interfaces (inlets/outlets) of PDMS
chips consisted of punched or molded holes are prone to
leakage and awkward to connect in comparison to the
leak-free connectors (e.g., Luer-lock, barbed connectors);
(3) engineering expertise and equipment (e.g., computer,
pressure sources, software) required for the operation of
fluidic valves and connection of chips are absent in most
biomedical laboratories; (4) multiple layers of PDMS must
be fabricated by standard methods that is tediously slow
and then sealed together for a 3D channel (multilevel
channels or a single channel with different sizes), which
limits the complexity of 3D constructs; (5) PDMS is a very
porous matrix that swells in organic solvents, resulting in
the loss of solvent into the microchannel walls,
detachment of the seal between the channels and the surface,
and alterations of the channel geometries.
Because PDMS is unable to be directly printed, other
materials are also employed. Biocompatible and
transparent resins provide the possibility of fabricating
biomedical devices by stereolithography (SL) although most
of SL resins are non-biocompatible and translucent or
opaque materials in the jewelry and structural modeling.
Other properties in choosing the resin for SL-fabricated
fluidic devices are gas permeability, hydrophobicity, and
chemical stability in the presence of solvents. The
currently popular biostable resins are based on polyester/
polyether oligomers with acrylate or methacrylate
functions and biodegradable composites of
methacrylatefunctionalized polyesters. Fibers of polyethylene and
nylon have proven to be an excellent choice for
preparation of 3D elements. Heating the polymer wires above
their glass-transition temperatures but below their
melting points allows for forming the desired shapes.
WaterShed is nearly colorless with a clarity, flexibility,
and hardness similar to polycarbonate or poly(methyl
methacrylate). Furthermore, it does not swell in water and
meets biocompatibility standards ISO 10993–5
(cytotoxicity), ISO 10993–10 (sensitization), ISO 10993–10
(irritation), and USP Class VI [
]. However, longer-term
cytocompatibility of WaterShed needs further investigation
]. Internal processing of PMMA, PDMS, polystyrene
(PS), and polyvinyl alcohol (PVA) polymers has also been
investigated. Polypropylene (PP) is an attractive material for
the fabrication of micro- and milli-reactionware as it is a
robust, flexible, and chemically inert polymer, and
significantly less expensive than PDMS.
Naturally derived polymer (e.g., alginate, gelatin,
collagen, fibrinogen, agarose, chitosan, fibrin, and hyaluronic
acid) or modified proteins (gelatin methacrylate) isolated
from animal or human tissue for 3D bioprinting has the
similar property to the human extracellular matrix
(ECM) and inherent bioactivity [
synthetic polymers and molecules [e.g., polyethylene
glycol (PEG), PEG amine] can be tailored to specific
physical properties for specific applications but has poor
biocompatibility, toxicity, and loss of mechanical
properties during degradation [
]. Synthetic hydrogels are
both hydrophilic and absorbent, especially attractive for
regenerative medicine [
]. Synthetic–natural mixtures
are also used to combine their advantages. In the 3D
bio-printing of vascularized tissue constructs, the
preparation of bioink composed of cells suspended in a
liquid pre-gel solution is critical [
]. During the
printing process using mechanical extrusion, the bioink is
gelled by polymer crosslinkers, photo activation, or
thermal activation to form a hydrogel that physically
constrains the homogeneously suspended intact cells
without compromising the cell viability and organelle
activity illustrated by fluorescent assays and organelle
tracking even after 48 h of culture, which is due to the
similar mechanical characteristics of 3D crosslinked
hydrophilic polymer networks in the hydrogels to that of
]. By varying the concentration of crosslink, the
hydrogel can be tuned to be “soft” or “robust” gels [
The cellular proliferation should be high to populate the
printed construct but be maintained at an appropriate
rate to achieve tissue homeostasis without hyperplasia.
The DNA bioink is advantageous over synthetic polymer
hydrogel because of its higher biodegradability.
Stereolithography is the most popular 3D printing
approach to directly print the micro-channels or create
modular structures. As opposed to molding processes,
SL is fully digital, amenable to finite element modeling
(FEM), intrinsically modular, and able to simplify the
commercialization pathway [
]. The single-photon
polymerization (1PP) process occurs near the surface of
a photosensitive resin. The outcome of SL is dependent
upon the laser spot size and the absorption spectra of
the photo-resins. It presents an inherent advantage in
the production of 3D structures over other lithographic
methods (e.g., photolithography and soft lithography)
owing to no need of alignment or bonding. Laser raster
scanning, laser vector scanning, and digital light
processing (DLP) have been developed for curing the resins
in commercial SL instruments [
]. In DLP-SL, an entire
layer of resin is exposed at once so that its resolution is
determined by the projected pixel size. Digital
micromirror display (DMD) technology and commercially
available projectors allow reducing the price of DLP-SL
printers significantly (e.g., ~$100). The structural fidelity
is superior in the free surface technique over the
constrained one because the mechanical separation in the
bat configuration can induce stress fractures or bend of
delicate features and increase roughness between layers.
However, the resin reservoir depth limits the object
height in the free surface technique, but not in the bat
configuration. Furthermore, the curing time is shorter in
the bat configuration, where the photo-polymerization
inhibited by oxygen occurs away from the air-resin
interface. The achievable layer thickness is only dependent
on the Z stage resolution, but not the resin viscosity.
The large discrepancy in the price of SL printers is
attributed to resolution, build area and speed. Printed
reusable templates have resolutions of 50 μm and up to
10 μm in localized hindrances, and can be fabricated
within 20 min at an average cost of $0.48 [
]. The main
limiting factors of SL are the effective drainage of the
uncured liquid resin, optical clarity, and Z-height
resolution. SL printers cannot change the printing materials
easily, but both resolution and surface finish are
sufficient to make PDMS templates with the combination of
thick and thin features.
In two-photon polymerization (2PP), two photons from
femtosecond pulsed near-infrared lasers are absorbed
simultaneously by the photo-initiator, directly recording or
writing an arbitrary polymeric 3D pattern into a volume of
photosensitive material. 2PP is not limited by the laser
diffraction, resulting in much higher resolution (e.g.,
~100 nm) in comparison to 1PP. For the negative
photoresists such as those containing acrylic oligomers or epoxy
resins, 2PP produces the crosslinking of polymer chains
through radical polymerization and makes the exposed area
insoluble in the solvent, which provides the possibility of
directly writing the structure. For the positive photoresists,
2PP causes the polymeric chains to break and become
soluble in the solvent to write the reverse structure. Although
commercial negative photoresists have better capacities of
modeling and conformity, they are not commonly used for
fabricating fluidic chips because of the long processing
time. Femtosecond lasers induce a local phase change in
the photo-sensitive glass (e.g., Foturan) from amorphous to
crystalline and can produce sub-wavelength features as
non-linear absorption is not limited by optical diffraction.
However, they are too expensive, 3–6 times as nanosecond
CO2, excimer, and Nd:YAG systems.
Selective laser melting and sintering (SLS)
This technique uses the powders with a high purity and
properties similar to those obtained by traditional
fabrication in the sintering so that it is advantageous over
other 3D printing techniques. SLS is also used to write
metal patterns onto polymers (e.g., PDMS), which has
great potential in the design of biosensors. A variety of
materials including metals, ceramics, and polymers,
which are typically proprietary with poorly characterized
surface properties, are used. Finer particles are used to
produce accurate and smooth parts, but difficult to
spread and handle. By contrast, larger particles facilitate
powder delivery and process but hinder surface finish,
resolution, and layer thickness. However, the obstacle of
fabricating fluidic devices by SLS is that it is very
difficult to remove the powder precursor from small cavities.
Fused deposition modeling (FDM)
FDM can print a large number of cheap and
biocompatible polymers, such as acrylonitrile butadiene styrene
(ABS), poly-lactic acid (PLA), polycarbonate,
polyethylene terephthalate (PET), polyamide, and polystyrene
owing to its advantages of safety, reliability, easiness in
the use, office friendness, low price, low levels of fumes
from polymer at high temperatures, and no requirement
of post-processing. FDM of liquid precursors, such as
metallic solutions, hydrogels, and cell-laden solutions,
has been implemented in the manufacture of LEDs,
batteries, strain gauges on flexible substrates, antennas,
interconnects, and electrodes in biological tissue.
However, the structural strength of FDM printed
structures is low and prone to compressive stress fracture
because the extruded material immediately hardens and
the adjacent layers are not well fused. There is a
tradeoff between printing resolution and surface finish, and
the smallest fluidic channel achievable (~100 μm) is still
larger than those made by SL.
Polyjet or multi-jet modeling (MJM)
MJM is attractive for fluidic applications because of high
resolution and capability of printing multi-materials
(over 100 different raw materials, including 22 from
Stratasys, 38 from 3D System, and many ones used in
the lab). Inkjet operates either in continuous or
dropon-demand (DoD) mode. The polyjet printer produces
smooth features with the surface roughness of 0.47 μm
in comparison to that of 42.97 μm in FDM. The inkjets
can deliver simultaneously multiple materials with a
wide range of properties (e.g., hard and soft plastics,
elastomers) in different colors. However, the currently
available materials are proprietary and expensive, and
rigorous biocompatibility and bio-functionality
investigations are required. The smallest printed fluidic channels
are approximately 200 μm [
]. Comparison of these 3D
printing technologies is listed in Table 1, and their
diagrams are shown in Fig. 1 [
]. Surface roughness
induced van der Waals, electrostatic, and steric forces are
unique to microfluidic flow. The induced shear stress
may cause transient pores on cell membrane for
complete cellular death and enhanced apoptosis . In
addition, silicone or mineral oil has been used to match
the refractive index of fabricated devices in order to
reduce the spherical aberrations for optical imaging.
Realistic 3D channels may not always be formed by
onestep molding because the molding material and the master
will be interlocked with each other. Peeling off the PDMS
from the master under its partially cured state will generate
a crack which will be self-closed afterward due to its
elasticity and self-adhesion by further thermal curing. Oxygen
plasma treatment followed by silanization is to coat a
monolayer of fluorinated molecules on the 3D printed
master to prevent PDMS from sticking to it. Heating at 130 °C
is able to remove the unreacted additives and monomers
inside the printed master. Such
“heating–plasma–silanization” strategy allows researchers to fabricate 3D fluidic
chips easily without bonding and alignment repeatedly and
the clean room. Print-and-peel (PAP) techniques, printing
the masters directly for casting polymers and adding 3D
components onto the masters for single-molding in the
bulky slabs, are facile and expedient in prototyping fluidic
devices with regular office equipment (see Fig. 2). Ink or
toner is deposited on the surface of the smooth and
nonabsorptive substrate (e.g., overhead transparency films)
leaving positive-relief printout features by LaserJet or solid
ink printer. PAP has been utilized to fabricate polymer
micromixers, capillary electrophoresis, valves, gradient
generators, optical waveguides, microelectrodes for μTAS.
While the channels fabricated by photolithography have
almost rectangular cross sections, those made from
LaserJet- and solid-ink-printed masters have trapezoidal and
round-bottom cross sections, respectively. The smallest
lateral feature size reproduced on printed masters is around
100 μm while the heights of the features do not exceed 15
μm. However, the durability of the printed masters needs
further improvement for mass production.
Although the fluidic devices printed by inkjet-printed
are inexpensive, they are limited to planar channels on
the glass surface. Alternatively, double helix channel is
also possible. Strands in such shape are first
manufactured and anchored, and their inner surfaces are
separated by 2 mm through the posts on both ends. The
PDMS prepolymer is then cast around the double helix
structure, and the mold is then manually extracted
following the curing process to create helical channels
(500–1000 μm). The mold material inhibits full curing
of the PDMS at the mold/PDMS interface [
However, such manual extraction method is only feasible
for certainly shaped scaffolds.
Most subtractive methods can only remove material
from the surfaces and are inappropriate for fabricating
complex fluidic chips. Materials such as carbohydrates,
hydrogel, metals or polymers are used as sacrificial
templates and removed from the solidified polymer. The
scaffold plastic polymer (e.g., ABS) is suspended into
liquid PDMS and dissolved using a PDMS-inert solvent
(e.g., acetone for 12 h) after curing the PDMS, leaving
an empty cavity inside the PDMS. The swelling ratio of
acetone for PDMS is as low as 1.06 [
Scaffoldremoval method is powerful and versatile in creating
multilevel and intricate fluidic channels. Integrating
external elements directly in the fluidic device is
desirable for LOC, but difficult to achieve using standard
PDMS fabrication methods. Heating coils, RF circuitry
or electronic components are also able to be embedded.
However, reliably clearing a sacrificial material from an
enclosed channel is limited by diffusion and quite
challenging to producing arbitrary microfluidic networks in
a single step. In addition, harsh condition, such as high
temperatures for creating or removing [
] and applying
heavy swelling for pulling out the template [
], are the
limitation of this approach.
Fused sugar has an advantage as a sacrificial template to
fabricate smooth channels owing to its efficiency and
]. Maltitol is selected due to it stable melt
status, suitable surface tension, and high water solubility.
The process is usually within 5 min. The diameter of
printed sugar filaments is affected by nozzle diameter, air
pressure, printing temperature and speed . PDMS is
then cast onto the sugar structures. Such process is
repeated for each layer. Finally, PDMS is solidified at 85 °C
for 25 min. Fluidic chips are immersed into hot water to
dissolve the sugar lines without further sealing (see Fig. 3).
Low cost is a significant advantage of sugar printer (~
$800). However, the structure without appropriate
supporting materials may restrict the design complexity because of
the structure collapse occurred between the large junction
space of different channels due to surface tension and
gravity of hot sugar filaments. Similarly, PDMS fluidic devices
can also be fabricated with 3D wax jetting by a glass nozzle
and a lead zirconate titanate (PZT) actuator [
One of the liquid metals in the use is EGaIn, a eutectic
alloy of gallium (Ga) and indium (In) in a 3:1 ratio by
weight with a melting point of 15.5 °C. Because of the
passivating oxide skin thin shell forms instantaneously
on the metal surface at room temperature. 3D printing
is done using a micropositioning stage and a pneumatic
air dispensing from a syringe. The printed structures are
small due to the short distance between the nozzle tip
and the substrate (<100 μm), and the movement of the
nozzle generates stresses that could “neck” the metal
filament spanning across this stand-off distance. Casting
and curing polymer onto the printed features define the
microchannel wall. After the printing, the EGaIn can be
withdrawn using electrochemistry (e.g., 1 M HCl)
because the drop and a bead of liquid metal at the other
end act as anode and cathode, respectively, which is less
harsh than acid. The metal bead also lowers the Laplace
pressure at the outlet, making it easier for withdrawing the
liquid metal (see Fig. 4). The height of structures printed is
limited by the stability of the oxide skin (e.g., ~4 mm for
EGaIn) especially in the embedding due to shear forces.
Direct ink writing (DIW) is an attractive method for
creating 3D microvascular structures [
]. A fugitive
organic ink is patterned into the desired motif,
encapsulated in a thermally or photocurable resin, and
subsequently removed by liquefaction to yield uniform
microchannels interconnected. Omnidirectional printing
(ODP) is a new variant of DIW and obviates the need
for layer patterning (see Fig. 5) [
]. The deposition
nozzle is inserted into a photocurable gel reservoir, which is
formed by pouring 25 w/w% Pluronic F127-diacrylate
into a silicone mold at 4°C and slowly solidifying at room
temperature and can physically support the patterned
features. Air pressure extrusion is applied to print the fugitive
ink filaments, whose size is linearly proportional to
applied pressures. As the deposition nozzle translates
through the reservoir during printing, void space is
generated locally and immediately filled by a 20 w/w% Pluronic
F127-diacrylate fluid on the top of the gel reservoir
because the liquid filler has identical chemical
functionality, but a significantly lower viscosity than the
photopolymerizable reservoir. Afterward, the gel reservoir and
fluid filler are solidified via photo-polymerization under
365 nm UV light for 5 min to form a mechanically robust,
chemically crosslinked matrix. Because the fugitive ink
filled in the printing nozzles from 10 μm to 200 μm
in diameter (an aqueous solution of Pluronic F127 in
23 w/w%) has not been chemically modified and
pronounced shear thinning behavior (shear modulus >10 kPa),
it can be removed by liquefaction below the critical
temperature (< 10°C) under a light vacuum to yield the
microchannel network. This approach allows
omnidirectional freeform fabrication of 3D biomimetic microvascular
networks composed of a hierarchical, 3-generation
branching topology with various diameters from 200
to 600 μm [
A large number of biomaterials (e.g., living cells and
growth factors) can be directly printed using a 3D
bioprinter. Fabrication of complex and heterogeneous
structures using multi-head systems is relatively slow, which
limits their use for cell-laden construct. Using integrated
bioprinting-fluidics technology, the flow of different
bioinks or even ECM components can be integrated into
fibers or droplets, which opens new routes for creating
realistic tissue fibers on demand. Robust hydrogels can
be extruded through the dispenser, but soft gels are in
the form of continuous polymer strands ideal for
building constructs. 3D cell encapsulation is advantageous
over conventional 2D cell culture in the cellular
morphology, proliferation, and gene and protein expression
because of improved cell–cell contacts and cell–matrix
]. Inkjet can print bioinks consisting of
cells, DNA, and biomaterials [
], while FDM is able
to create 3D multi-material scaffolds for cell seeding
. Mechanical forces during the bio-printing are
determined by extrusion speed, nozzle diameter, the
viscosity of hydrogels, and temperature. The shear forces
applied on embedded cells increase with the decrease of
nozzle diameter and the increase of the extrusion speed
and the viscosity of hydrogels by decreasing the chamber
or nozzle temperature [
]. Simultaneous extrusion of
an alginate and a calcium ion solution through the inner
and outer needles, respectively, of the coaxial extruder
permits the formation of a gel fiber at the tip and lays it
according to the design. The optimal formulation is 4%
w/v alginate and a 0.3 M solution of CaCl2. Such method
obtains macroscopic and porous 3D structures with
single fiber thickness in the order of 100 μm. However, high
concentrations of crosslink materials (>2%) have a
negative impact on the cell’s salt balance. The use of two
independently cross-linkable hydrogels allows the tuning
of mechanical properties of the cell-laden fibers to
mimic the morphological and mechanical features of
native tissue. The concentration and densities of two
different hydrogel-precursor polymers are adjusted for a
higher printing resolution of cell-laden fibers and the ideal
microenvironment for cell spreading and organization.
Overall, the cells should be robust for physical forces (e.g.,
shear stress and pressure) and biological stressors (e.g.,
the presence of toxins, enzymes, and nonphysiological
pH) in the bioprinting. Current 3D bioprinting approaches
involve biomimicry, autonomous self-assembly, and
minitissue building blocks [
Paper-based fluidic device
Paper-based fluidic devices are recently developed and
have significance in the simple fabrication for mass
production, low cost, ease of transportation, storage, and
disposition, simple liquid motion without excessive
equipment, but the high efficiency. Fabricating a micro
paper-based analytical device (μPAD) by wax printing
involves only two main procedures: printing wax
patterns on the paper and then melting the wax across the
paper to form hydrophobic barriers both laterally and
vertically to prevent the fluid mixing passing through
the device. The low production cost and complexity
levels in the manufacture of μPAD are appropriate for
prototyping at a large scale. Among the numerous
techniques utilized to create channels on hydrophilic paper,
solid ink (wax) printers are the most promising one for
smooth features [
] while the granular structure of the
LaserJet toners is conspicuous on the replicas. The
relatively low melting point of the wax prevents the cast
PDMS from curing at elevated temperature. The lateral
dimensions achievable with office-grade solid-ink printers
are about 200–300 μm while that of office-grade LaserJet
is 50 μm [
]. Although the features are not accurate and
sharp at the edges, they are sufficient in detecting
substances due to color change in the test assay. Overall,
μPAD may develop to rapid, cheap, flexible, and reliable
devices for clinical emergency and large-scale use [
This fabrication method is quite new, and the control over
flow rates, mixing, and interaction times between sample
and reagents needs to be improved.
Microfluidic systems are valuable tools in flow cytometry,
cellular assays (e.g., cytotoxicity or cellular stress assays),
cell sorting, manipulation, and imaging, molecular
analysis, cell response to chemical and physical stimuli and
tissue engineering because of precise control of small
volumes of fluids over short distances. Array design provides
the possibility of parallel measurement of a large number
of samples. Some of the emerging applications of fluidic
channels in biology and biomedical engineering are listed
below to illustrate their technical advantages. These
applications are summarized and listed in Table 2.
Molecule and protein detection
A variety of electrode materials (e.g., carbon, platinum,
gold, silver) can be easily integrated into microfluidic
devices for various applications (e.g., neurotransmitter
detection, NO measurement, oxygen tension in a stream
of red blood cells) along with other functionalities (e.g.,
fluidic interconnects and membrane inserts) for
molecule analysis (e.g., ATP via chemiluminescence).
Essentially, the electrode fitting is removable and reusable. 3D
printed fluidic channels not only change the strategy of
research collaboration but also the perceived limitations
of the biological experiment, where the spatial control of
samples or cells is critical [
]. Biosensors could also be
integrated and placed consecutively in the fluidic chips.
Paper-based fluidic devices are biodegradable, cost
effective in the disease diagnosis, and easy in almost all
environments. The amounts of glucose and protein in
the paper fluidic channels are proportional to the color
change of each assay [
]. The smallest width of the
printed hydrophobic barrier is 400 μm, and expanded to
1000 μm after melting.
cell deposition & simulation
3D tissue constructs
organ conformal biopsy
Milli- & micro-fluidic
various electrode materials integrated into
microchannel; various measurement and
functionalities; easy operation in all
at the resolution of individual cells, the
possible molecular interactions between
cells, 3D concentration of gradients, precise
control of fluids, reduced reagent/sample
consumption, robust and automated
multiple population, no external force
required, little physical damage to cells
precise control over various cellular
microenvironment, easy formation of
desired structures, high throughput,
reproducible, multi-layer structures
rich diagnostic information, continuous
monitoring, direct coupling
rapid production and design optimization,
quick and versatile material synthesis, high
accurate position of various tissue samples
fabrication complexity with
increased number of embedded
sensors, large fabrication error
width in paper channel
rather large fluidic channels,
discrepancy between the
printed and designed
hard to predict and control the
flow behavior in a channel with
varying curvatures, small bacterial
concentration in the detection
trade-off between cell density of
bioink and nozzle size
throughput is limited by large
components with intricate
unknown long-term effects for
low output volume, inability at high
pressure and temperature
Cell deposition and simulation
The complicated topology of the 3D microfluidic
network in the stamp makes it versatile to pattern
multiple proteins and cells. However, the cells have
limited migration and growth and stop dividing once
they form a confluent layer between the stamp and
the substrate [
], and continue to spread and divide
once the stamp is removed. There are two levels of
membrane structures: one as the channel plane open
for contact with the substrate, and the other as vias
connecting channels in the membrane to those in
the slab. Autoclaving is necessary to improve the
viability of cells inside the PDMS stamp [
deposition of multiple cell types and proteins in
complex, discontinuous, well-defined patterns has
great value because in vivo ECVs (tumor cells) attract
and direct the growth of BCEs (capillary cells) for
tumor angiogenesis and nutrients and oxygen supply. It
makes 3D micromolding in capillaries (MIMIC) a
powerful technique in investigating the differential and
competitive attraction of capillary endothelial cells to different
tumor cells, which can be developed to a simple, standard,
and quantitative in vitro assay for evaluating the
angiogenic potential (see Fig. 6). It can also be used in
investigating the functional significance of tissue architecture at
the resolution of individual cells, and the molecular
interactions between cells that underlie processes of
embryonic morphogenesis and formation of the blood–brain
Concentration gradients of soluble factors (e.g., growth
factors, chemokines, and gas molecules) are essential for
physiological and pathological processes in vivo. Thus,
the generation of 3D concentration gradients has strong
implications for tissue engineering and drug screening.
There is a gradient of physical properties from central to
the peripheral vascular tree. For example, the arteries
closer to the heart are thicker and more compliant
whereas arteries further along the vascular tree are
considerably thinner and stiffer. Microfluidic technologies
provide benefits over conventional cell culture and
experimental systems because of the precise control of
fluids (on the scale of fl and nl), cost effectiveness,
reduced reagent/sample consumption, and robustness
via automated experimental procedures [
mold fabricated using the 3D printing from a single
material shows superior mechanical stability in comparison
to photoresists on silicon. However, current fluidic
channels are still quite large. The discrepancy between the
printed and designed one is 30–70 μm at the channel
height of 250 μm and increases significantly at the
channel height of 100 μm . Those below 50 μm are not
Microscopic printing also enables to organize multiple
populations of bacteria within 3D geometry (e.g.,
adjacent, nested, and free-floating colonies). To investigate
the behavior of small microbial aggregates (e.g., whole
bacteria, cells, ATP, oxygen, and other essential
biomolecules), a number of microfabrication technologies have
been developed to confine bacteria within the fluidic
devices, cavities, and liquid droplets for assaying antibiotic
resistance and enzymatic activity. Such process of
encapsulating cells often restricts mass transport, which is
incompatible with growth and signaling between
physically isolated populations. The resistance of one
pathogenic species to an antibiotic can enhance the
resistance of a second species by virtue of their 3D
relationship. Moreover, this fabrication approach using
bioprinting can define bacterial micro-colonies in animal
hosts for infections development in vivo [
]. 3D printed
chip by FDM is also suitable for bacterial cultivation,
DNA isolation, PCR, and detection of an amplified gene
using gold nanoparticle (AuNP) probes for early
diagnosis with compactness and low cost.
Because of fast proliferation, sensitive measurement of
bacteria at the early stage is critical for preventing
foodborne diseases [
]. Microbial cultivation-based
detection is accurate and reliable as a golden standard
method. However, its application is limited to laboratory
measurements owing to the intensive consumption of
time and labor. Size-based separation techniques are
preferred because of no requirement of a complicated
labeling procedure. The inertial focusing by Dean drag
force has been successfully used to separate cells and
particles in a 2D PDMS substrate with easy control of
the operation condition, but no external force required
and little physical damage to the cells. However, varying
the curvature radius in a spiral channel with different
Dean numbers makes the flow behavior difficult to
predict and control. A helical microchannel around a
cylindrical chamber was fabricated using stereolithography in
a compact size with a constant radius of curvature (see
Fig. 7). Large antibody-functionalized magnetic
nanoparticle (Fe3O4) clusters are focused near the inner wall of
the microchannel where Dean drag force and the
magnetic lift force proportional to the particle volume
are balanced. To improve the separation, a sheath flow is
introduced to push the particles to the outer wall of the
microchannel and help their trapping in the strong Dean
vortex cores. The detection limit is 10 cfu/mL for E. coli
bacteria in buffer samples and 100 cfu/mL in milk due to the
less capture efficiency by the presence of interferents [
3D tissue constructs
Biological tissues and organs have a large number of
microvascular networks facilitating oxygen and nutrient
delivery and waste removal from the surrounding cells.
Mimicking such network is of considerable importance
for self-healing [
], replacing a damaged native
blood vessel, tissue engineering , organ printing [
], tubulogenesis and vascular
], cardiovascular pathology,
pharmacological modeling, drug testing (e.g., functionalization
of biomaterials with proangiogenic agents), and
biomedical devices. For example, skin mimicking samples with
healing agents in the synthetic microvascular networks
could repair the damage repeatedly [
techniques have emerged to induce the formation of
vascular structure within tissues and can be classified
into either pre-vascularization-based or
vasculogenesisand angiogenesis-based types. Due to a lack of effective
vascularization, there are severe limitations in the
clinical development of vascularized complex 3D tissues,
particularly those large vital organs (e.g., liver, kidney,
and heart) [
]. While the pre-vascularization techniques
provide readily available channels for immediate
perfusion of growth media or blood and fabrication of larger
blood vessels, they are not suitable for vascular capillary
beds with cascading bifurcations down to a few micron
sizes. The vasculogenesis- and angiogenesis-based
approaches, on the other hand, provide very limited
control over the temporal and spatial factors, require
days to weeks before cells can organize and grow
perfusable lumens, and are not suitable for formation of
vascular structures for suturing and anastomosis with
the host vasculatures. The combination of hydrogels,
microfabrication techniques, and microfluidic systems
may overcome the challenges of developing an artificial
] by offering precise control
over various cellular microenvironment including fluid
flow, chemical gradients, localized ECM as well as the
microenvironmental cues such as mechanical properties
(e.g., stiffness), chemical properties (e.g., ligand density
and orientation), and topographic features (e.g., different
cell substrate affinity). The vascular channel is constantly
perfused and resides on a biologically relevant and
porous matrix where other cells can be introduced easily to
form the desired structures. The embedded live cells and
growth factors along with biomaterial channels at
precisely controlled locations to mimic the native tissue
architecture can facilitate the delivery of nutrient-laden
fluids and promote cell viability [
]. It has a great
potential in tissue engineering because various
functional tissues can be fabricated with appropriate
structures and cell compositions (e.g., EC, SMC,
fibroblasts, progenitor cells or various stem cells) in various
sizes, high throughput, and reproducible fashion .
However, the fabrication of microvascular networks
composed of complex, hierarchical 3D architectures is
still quite challenging. Building appropriate vascular
structure is critical to vitalize thick tissue which has
difficulties in survival and proliferation due to diffusion
limitation over a few hundred micrometers [
Culturing under dynamic conditions could sustain cell
viability deep inside the scaffold. During the incubation
process, HUVECs capture within the gelatin sink down
slowly and attach to the inner surface of the channel
]. HUVECs cover 70–80% of the inner surface area of
the fluidic channel on Day 0. The cells proliferate and
cover the entire inner surface within 2–3 days. Under
the flow culture condition, HUVECs on the channel wall
are elongated and aligned along the flow direction over
time. In comparison, HUVECs cultured in the static
condition escape from the channel edge, actively
invading into the collagen scaffold and forming angiogenic
sprouts (see Fig. 8). The sprouting initiates on Days 3–4
all over on the channel wall and extends up to 400 μm
on Day 7. As the sprouts continue to invade into the
collagen matrix, they become longer, contain
progressively more cells, and begin to branch. Stereotypical
sprouting morphology is observed in these sprouts,
presenting thin filopodia-like protrusions at the sprout tips.
Frequently, cell migration into the collagen matrix
occurs exclusively by angiogenic invasion and sprouting.
Among the most extensively investigated methods for the
effective introduction of angiogenesis in vivo is the control
and regulation of the spatial and temporal distribution of
common angiogenic growth factors (GFs), such as
concentration gradients of VEGF, in a cell-laden or cell-seeded
hydrogel in a microfluidic device. The endothelial cells
(ECs) tend to migrate from the region in low-GF
concentration toward that in high GF concentration, thereby
aligning themselves into well-organized structures and
enhancing the capillary-like tubular structure formation
]. Hollow, calcium-polymerized alginate tubes can be
easily patterned using 3D printing techniques. Its diameter
can be precisely controlled in the range of 500–2000 μm by
changing the flow rates or nozzle speed. The structural
rigidity of these constructs allows the fabrication of
multilayered structures and maintenance of their hollow form
without causing the collapse in lower layers [
The development of highly organized and functional 3D
tissue constructs is still challenging. Precisely positioning
different types of cells and biomaterials to resemble the in
vivo environments is a major problem despite significant
advances in 3D bio-printing [
biomaterials are required to enable spreading and migration of the
cells, printing low-viscosity bioinks for the use of a smaller
nozzle and faster dispensing speed for a higher printing
resolution and shorter fabrication time, and fast gelation
process to support the cells located inside and outside of
the printed scaffold and to create thick constructs with
high cell viability. High-density cells (107 cells/mL) within
the polymeric solution can reduce the possible shear stress
applied to the cells during the bio-printing. Furthermore,
the fluidic platform incorporated to a 3D printing system
can rapidly deposit multiple materials through an
extrusion system and precisely switch between different bioinks
and patterns, which allows the creation of heterogeneous
3D structures in the improved resolution and efficiency
(see Fig. 9). For example, low concentrations of GelMA
hydrogels (<5% w/v) with a low degree of acryloyl
modification show spontaneous organization of
encapsulated cells, such as human mesenchymal stem cells
(hMSC) and ECs, in comparison to the high
concentration (>10% w/v GelMA) with a high degree of
acryloyl modification [
There are significant differences in cell behavior
between 2D and 3D models in protein expression and
gradients, drug response, cell migration, morphology,
proliferation, and viability [
]. Most 2D cells float from
the culture dish while 3D cell spheroids in the hydrogel
are still maintained within the constructs. The metabolic
activity in 3D and 2D Hela culture with the addition of
paclitaxel is 0.47 and 0.06 times, respectively, in
comparison to the control. It shows that 3D cell/hydrogel
constructs are important in supporting the long-term
proliferation of a large number of cells. Hela cells in the
3D model show higher MMP protein expression and
chemo-resistance than those in the 2D culture [
differences in gene expression patterns of 2D vs. 3D are
likely due to the matrix composition and stiffness that
ECs reside on (collagen vs. plastic surface). The
heterogeneous distribution of biological-relevant proteins and
growth factors in the tissue are essential for cell
signaling, proliferation, and migration. Bioprinting can
mimic the structure and function of in vivo cancer with
3D complexity (e.g., tumor heterogeneity, leaky and
poorly organized vascular structure) in a biomimetric
microenvironment in a high-throughput and
reproducible manner at low cost. Stromal cells co-printed with
tumor cells can naturally secrete ECM, growth factors,
and hormones so that structural differences between the
proteins used and the varying composition and materials
in the exogenous scaffolds can be avoided.
Histomorphological analysis showed adipose, stromal, epithelial,
and carcinoma compartmentalization in the printed
cancer models [
]. Living microarchitectures bioprinted
from human cells are more realistic for creating cancer
models without the cross-species difference which leads
to the inaccurate prediction of the animal models in
human testing [
Organs on a chip are microengineered tissues cultured in
specifically designed and fabricated microfluidic
bioreactors, supplying oxygen, nutrients, and growth factors and
removing waste, to mimic the structure of human tissues
better than current models. Because of the translational
challenges associated with 2D monolayer cultures (e.g.,
inability of replicating higher order features and
trajectories of human body) as well as ethical and economic
concerns of small animal experiment, 3D printing of the
next generation of microphysiological neural
systems-ona-chip (NSCs) can model human neurological diseases
using stem cells in the construction of patient-specific
NSCs, develop personalized neurology, and facilitate the
preclinical drug screening with more flexibility,
robustness, and efficiency in controlling and monitoring system
]. 3D printing affords repeatability and
robustness in multi-layer and diverse-material (e.g.,
conductive materials as the substrate for stimulation and
monitoring) fabrication for interweaving biology with
scaffold and functional materials directly using the anatomical
geometry obtained from the medical diagnosis.
Microextrusion is more popular in developing NSCs because it
is compatible with cell suspensions, cell-laden hydrogel,
and thermoplastics. Microfluidic NSCs can model neurite
outgrowth, fluid handling for perfusion, convective flow of
nutrients and biochemical cues (e.g., cell migration, cell
signaling, and gene expression), and mechanical actuation
(e.g., shear stress and dynamic scaffold deformation).
Laminar flow inside it allows the generation of complex
and highly controllable fluid flow regimes. The bioreactor
can maintain the viability of tissue constructs and
accelerate tissue fusion, remodeling, and maturation. However,
the throughput is limited by large components with
intricate geometries. In compartmentalized NSCs, the
coculture of multiple cell types, media, and biochemical cues
allows the investigation of cell-cell interaction in tissue
self-assembly and circuit mapping (see Fig. 10). In
comparison, hydrogel NSCs have more flexibility to design
heterogeneous tissue by including multiple cell types,
crosslinking hydrogels with different ECM compositions
or different cell-laden hydrogels.
Another important platform is liver-on-a-chip for
longterm culture (e.g., 30 days) of 3D human HepG2/C3A
spheroids encapsulated with GelMA hydrogel for drug
toxicity assessment in a bioreactor with continuous
perfusion and in situ monitoring of the cellular functionality by
analyzing the concentration of secreted biomarkers [
Culturing the cell spheroids could enhance homotypic
cell-cell interactions with aggregated hepatocytes and
improve functional outcomes. The toxic response in such
hepatic construct was found similar to that of animal and
in vitro models. Further investigation is required to yield
cells with phenotype and functions similar to those of
mature hepatocytes (e.g., drug metabolism, bile formation,
and production of blood clotting factors and glucose). The
other developed and successful models are for the
bloodbrain barrier, lung, intestinal [
Organ conformal biopsy
Conformal microfluidic devices work as a minimally
invasive “biopsy” for isolation and profiling of biomarkers
from whole organs within a clinically relevant interval. 3D
printing exhibits compatbility with conformal
manufacturing of next-generation microfluidic devices and medical
imaging technology for whole organ healthcare (e.g., organ
assessment) because of a major advance in microfluidics
and the direct coupling of the device to the surfaces of
whole organs. The samples continuously isolated by the
3D printed organ-conforming microfluidic device provide
rich diagnostic information, such as biomarkers of
ischemic pathophysiology and metabolic activity. This
achievement could shift the paradigm for whole organ
preservation and assessment, thereby relieving the
organ shortage crisis through increased availability
and quality of donor organs [
Milli- and micro-fluidic reactionware
The rapid realization of configurable and scalable
reactors is highly desired in chemistry. The high surface
area-to-volume ratio and precise control of the reaction
environment are critical for LOC, but 3D printing
technologies have been overlooked due to a perceived
limitation of resolution. Reactionware is classified as
nano(1–100 nm), micro- (100 nm to 1 mm) or milli-fluidic
(1–10 mm) devices according to the dimension of
]. Due to the ease of micro- and
millireactionware fabrication, 3D printing facilitates rapid
production turn-around, iteration and optimization of
design based on experimental data at low cost once
faults or errors using composite catalyst-silicone
materials are found. Thus, it is straightforward to evolve the
design of milli-fluidic channels in terms of geometry,
inlets, outlets, and sizes, print in an appropriate material,
and perform organic, inorganic or materials syntheses in
one day, which allows versatility in the design and use of
specific reactionware for experimental users [
chemical reactions, such as organic synthesis of an amine
by two-step reductive amination and subsequent alkylation
of the secondary amine, the inorganic synthesis of large
polyoxometalate clusters, and the controlled synthesis of
gold nanoparticles, have been efficiently carried out in 3D
printed reactionware devices [
]. Nearly monodisperse
silver nanoparticles have been synthesized employing
miniaturized continuous flow oscillatory baffled reactors
(mCOBR) employing additive manufacturing with higher
temporal stability and superior control over particle size
distribution than tubular flow reactors [
In order to push the commercialization of microfluidic
channels, “killer apps” are necessary despite great
potentials shown in various applications. Specific tasks,
including the standard user interfaces, simple control on
the microfluidic systems, and commercial manufacture,
should be resolved. 3D printing has attracted attention
in fabricating fluidic networks due to its automation,
assembly-free, low costs, and continuously improved
resolution and throughput. Microfluidic channels have
great potentials for LOC (e.g., chaotic mixers, reagent
and buffer reservoirs, fluid homogenizers). But they have
limitations in terms of hardware, resolution, large
channels size, resin versatility, overall device dimensions, the
lack of control over resin formulation, subsequent
surface and bulk chemistry, and prototyping system cost
for uptake by “skill-less” biologists. With more
commercial microchannel products, “killer apps”, maybe in the
cell/protein detection and drug screening, will finally
show up in a few years. Although 3D printing cannot
substitute injection-molding at the mass production, it
can produce small batches (from single to hundreds of
parts) economically, efficiently, and environmentally
(minimum waste and no tooling) for a smooth transition
to injection-molding and easy design evolution,
permitting a “fail fast and often” strategy in the device
development based on the early and rapid empirical feedback
]. 3D printing will make most PDMS and plastic
molding in research laboratories but cannot completely
replace the photolithography. Printed fluidic devices can
dramatically reduce the barrier of sophisticated designs
and positively disrupt the developmental cycles [
Although the current resolution of 3D–printers does not
match that of soft lithography, 3D printing provides a
new route of integrating user interfaces and embedded
controls. The ability to clear the uncured or
partiallycured resin (in SL) or sacrificial polymers (in MJM) from
the 3D printed channels is important for the fabrication
resolution. However, the development in desktop SL and
MJM devices, photo-resins, multi-material 3D printing
as well as the expiration of patents and the emergence
of competing platforms are ushering in significantly
improved resolution, throughput, and functionality [
Systems with <10 μm resolution for <$10,000 are not
far-fetched; for example, a current system of $5000 can
achieve resolution of 25 μm [
]. Models up to 43
mm×27 mm×180 mm at the speeds of 20 mm/h in the
height were fabricated using a commercial 3D printer
costing $2300 with 500 mL of the clear resin of $138 for
any design complexity [
]. Oxygen inhibition of free
radical polymerization in air results in incomplete cure
and surface tackiness and is a widely encountered
obstacle. Continuous liquid interface production (CLIP) is
enabled by creating an oxygen-containing thin uncured
liquid layer between an oxygen-permeable window and
the cured part surface with the thickness of tens of μm
by judiciously selecting the photon flux and resin optical
and curing properties [
]. Subsequently, the resin
polymerization speed could be increased to hundreds of
mm/h. Plenty of work has been carried out in the open
fluidic channels, which is harder than printing templates,
but easier than printing enclosed channels by removing
the uncrosslinked resin. The use of both additive and
subtractive methods could lead to new devices
unfeasible with a one-mode approach, although material
incompatibilities between modes may be difficult in
practice. Another new trend is composite printing for
new functions, such as hybrid microfluidic/electronic
Fluid handling is a ubiquitous and tedious operation,
such as cell culture media in the benchtop research and
bodily fluids in clinical diagnostics. Usually, fluids are
transferred between containers by pipettors (prone to
operator’s error) or expensive robotic dispensers. Pumps,
valves, and mixers are critical for the fluids manipulation
and automation to reduce labor costs, speed up
processing, and enable mass parallelization. PDMS valves always
outperform plastic valves of similar size. The invention
of PDMS micro-valves and pumps revolutionizes the
microfluidics and heralds the miniaturization and
automation of multiple biomedical assays. The printed valves
can also work as functional modules. For example, two
valves in pair act as a switch while three valves in series
as a peristaltic pump. The absence of standardization in
interfacing PDMS devices with the peripherals is a major
bottleneck in the widespread adoption of LOC
technologies since the inlet/outlet connectors and tubing are
usually the most unreliable components in the channels.
SL-printed plastic 3D circuits with packaged connectors
can be built as interlocking modules that represent
existing industrial standards and are easy to operate.
The introduction of modular design paradigms and
integration of fluidic devices will amplify the efforts
of individual teams for industrial success. “Plug and
play” complex 3D milli-fluidic devices using flow
control, inter-connectable modular devices, and both passive
and active components allow the mixing, monitoring of
reaction and cell culture progress [
]. A sample library
of standardized components and connectors with
validated flow characteristics has been established, which
would allow the design and assemble complex 3D
microfluidic circuit as easily as that in the electronics industry
]. In addition, the modular design can also allow the
access to interior surfaces of microchannels to improve
the optical transparency using either mechanical or
chemical processes [
] so that analysis on these resulting
chips could be more accurate and reliable although
transparent 3D printing materials now have limited availability.
Fluidic devices with active valves and pumps as small as
10% of the volume and up to 1 million actuation could be
manufactured by DLP-SL and used for serial multiplexer
and mixer [
]. User-friendly fluid automation devices in
transparent and biocompatible channels can be printed by
non-engineers and integrated with other microfluidic
devices as the replacement for costly robotic pipettors or
tedious manual pipetting. Printing these devices requires
the digital file of various modules for new device assembly
and reconstruction with expanded functionality as well as
recyclability and electronic access to a printer by
nonexpert users without facility limitations. The combination
of rubber O-rings and metal pins improve module
connectivity. The inserted O-rings perfectly seal the
interconnections between the modules firmly and prevent leakage.
Because the predicted performance of a complex
multi-layer PDMS device from the ideal design is usually
different from the real performance, drastically reducing
the number of fabrication iterations in the development
of a complex device will save time and resources
significantly. The ability to fabricate a complex microfluidic
device in a single step has obvious advantages but challenges
]. Despite the enthusiasm of the early uptakers, its
applicability is limited partially by the technical inability to
print microfluidic channels reliably with dimensions less
than several hundred microns. MakerBot has created a
very vibrant website (“Thingiverse”) for sharing some
CAD designs with non-commercial (Creative Commons)
licenses. In addition, 3DSkema will soon launch an online
marketplace where designers can sell their licensed
designs. Web-based 3D–printing services are becoming
popular for designers owing to no requirement of
expensive molds and “minimum quantity” limit in the
Currently, photopolymer resins are used in 3D
printing technologies to make fluidic devices. New resins
exhibiting improved optical transparency, gas
permeability, and biocompatibility are continuously available,
which will favor further applications of 3D printing in
fluidic-based biological systems with optical
measurement. Optical transparency allows on-chip detection
while electrical insulation allows electrophoretic
separation. In addition, thermoplastics and elastomers are
used in non-photocurable techniques while soft
hydrogels in bioprinters. Cell viability is the key parameter in
the cytotoxicity tests of chemicals, cellular stress assays,
DNA sorting, single-cell behavior and cell manipulation,
especially organ-on-a-chip devices. The biocompatibility
and bio-functionality of the available 3D printing
materials are serious concerns but would facilitate cell
attachment on the printed surface [
]. Free radical is a
common concern associated with photo-polymerizable
hydrogel in tissue engineering, where photo-initiator
concentration compromises cell viability. Some of the
most promising resins from a biomedical perspective
(e.g., PEG-DA) are inexpensive and patent-free as they
have been used as biomaterials (photo-cross-linkable
hydrogels) for cell encapsulation for a long time. Although
initial biosafety and biocompatibility studies of 3D–
printed devices are encouraging (even implantable and
bioresorbable), longer-term in vitro cytotoxicity and in
vivo implant compatibility studies are greatly required.
Zebrafish embryos cultured in 3D–printed structures
made of Visijet crystal or Watershed showed
developmental defects, but no behavioral abnormalities were found
among those grown in leachate extracts of ABS and PLA.
Gradient generators, droplet extractors, and
isotachophoresis chips are successfully generated for the future
low-cost analytical applications. The ability to extract,
purify, label, or separate the sample within the device helps to
reduce analysis time and improve throughput. Samples
are then sensed and detected using optical (e.g.,
fluorescence), electrochemical (e.g., conductivity, amperometry,
and potentiometry), mass spectrometry, or biosensors.
New microfluidic designs integrating electrodes and
membrane inserts are successfully employed in the
electrochemical detection of neurotransmitters and viruses, the
collection of biologically relevant analysts (e.g., ATP), and
drug transportation [
Humidity, variance in the gelatin viscosity, dispenser
condition, and the printing parameters (e.g., air pressure,
valve opening time, and droplet spacing) can also
influence the channel diameter and structures. Air pressure
and valve opening time mainly influence the channel
width whereas the sequential number of printing has a
higher impact on the channel height. Increased channel
width results in the decreased migration speed of the
cells so that cancer cells move faster in smaller veins
than in large arteries. In addition, the proliferation of
HUVECs is suppressed under the flow condition, which
corresponds with previous studies that a long-term
exposure of the endothelium to shear inhibits cell
proliferation and reduces the metabolic rate [
culture media were consistently supplied, there was a
limitation on maintaining cell viability of vascular channel
probably because of high oxygen/nutrient consumption
when a large number of cells was embedded nearby the
channel. In addition, the cell viability decreases at the
higher concentration of gelatin encapsulation or the long
incubation time for gelatin liquefaction. One of the future
directions of 3D bio-printing is to create implantable thick
vascularized tissue constructs that could serve as artificial
organs or aid in their repair and regeneration.
Advances in 3D cell printing technology have enabled
the direct assembly of cells and extracellular matrix
materials to form in vitro cellular models for biology (e.g.,
hypoxia, tissue repair), the evolution of disease
pathogenesis and new drug discovery in printed 3D tumor
models in vitro [
]. Construction of a multi-layered
transport/μ-vascular channel network in a range of 100–
300 μm has a resolution of tens of microns. Although
the most effective way in tumor studies and anti-cancer
drug screening is in clinical trials, ethical and safety
limitations prevent it from wide acceptance. To overcome
this hurdle, preclinical tumor models are used to mimic
physiological tumorgenesis environments [
However, immunocompromised mice in use may show false
effects on tumor development and progression [
Constructing 3D microstructures can provide a virtual
environment that mimics the physical condition in vivo
appropriate for the growth of cells or micro-organisms
to a large extent and allow experiments to be conducted
with a more clinical or biological relevance compared to
culture in a Petri dish or flask [
]. Using this
technology, the composition of vascular cells and supporting
cells, flow rate/flow pattern, injection of soluble factors/
small molecules, and other media components could be
altered easily. Thus, vascular channels can be served as
an experimental model for diverse vascular
diseaserelated studies, for example, inflammation, immune
responses, and tumor angiogenesis [
]. The lack of a
simple and effective method to integrate vascular
network with engineered scaffolds and tissue constructs is
one of the greatest challenges in 3D tissue engineering
currently. Translation of 2D fabrication methods into a
3D complex vascularized tissue construct, integration of
channels across layers in the dimension of millimeters
and its mechanical stability should be paid attention
]. Neural stem cells can also be embedded along with
vasculatures and growth factors to examine their mutual
effect on network formation. The maturation process of
the engineered vasculature, capillary formation, and
angiogenic sprouting provide insight into the EC
behavior under 3D flow conditions for the investigation of
vascular biology. 3D cell printing has been reported in
the printing of in vitro liver tissues [
], adipose tissues
], bone tissues [
] and hybrid tissue constructs
with vascular-like networks [
Technical challenges in this field include, but are not
limited to the requirement of increased resolution,
speed, and compatibility with biologically relevant
]. The speed of fabrication should be increased
for constructs in clinically relevant sizes. Rapid
improvements in 3D printing resolution, even in low-cost
consumer-grade desktop 3D printers, are highly possible
for the rapid prototyping and cost-effective fabrication
in high resolution and therefore more precise control of
fluid flow [
]. The ability to image, map, and reproduce
complex 3D structures composed of biologically relevant
ECM proteins would be a major advancement for the
applications. Microenvironment for long-term cell
culture and growth in a user-friendly, highly-controllable,
and broadly-accessible manner would advance the
applicability of 3D printing to engineering physiological systems
]. The viability of encapsulated cells is impacted by
the processing time of the pre-gel bio-ink preparation and
the construct deposition and the sensitivity of different
cell types to external stresses. The use of multi-nozzle
printheads (e.g., photopolymer, UV, microplasma, bioink,
] for co-printing cell-laden hydrogels could
decrease the printing time and maximize the cell viability
]. It can be dedicated to printing the tissue
constructs and the microfluidic channels simultaneously. The
intersection of 3D printing for microfluidic fabrication
and bioprinting 3D tissues shows great promise to
organon-a-chip in single-step. With a combinatorial single-step
fabrication, organ-on-a-chip would be far more accessible
and cost-effective due to faster design iterations and
shorter turnaround times. Instrumented cardiac
microphysiological devices via multimaterial 3D printing has
been fabricated to monitor drug responses and
noninvasively measure tissue contractile stresses inside human
stem cell-derived laminar cardiac cell via embedded
sensors over 4 weeks . Cell reprogramming and directed
differentiation may provide high proliferation,
functionality, nonimmunogenity and robust cell populations [
Combination of various mature and multipotent cells can
be applied to efficiently reproduce the cell phenotypes for
specific tissue targets. Small embryoid bodies (EBs) are
more likely to have cardiomyocyte differentiation toward
ectoderm while larger EBs towards endoderm and
mesoderm. Spontaneous aggregation of inducible stem cells
and EBs leads to the inhomogeneous size distribution of
EBs and unpredictability in lineage differentiation. Thus,
control of EBs in uniform sizes and shapes is beneficial in
tissue engineering and regenerations [
]. 3D bioprinter
can be further integrated with minimally invasive surgical
robots to improve the healing procedure of the tissues
removed by the surgical intervention. With the use of
induced pluripotent stem cells derived from patients and
differentiated into particular cell types, it will eventually
be possible to generate organs-on-a-chip with a patient’s
own cells as personalized medical treatment. Furthermore,
a body-on-a-chip with multiple organs organized on a
single chip to better model the multiorgan interactions in
vivo is one of the investigative directions.
Simulation coupled with experiment can help in
understanding the effects of printing parameters on
cell viability [
]. Simulation can predict cell fate and
provide more parametric control over 3D cancer
models as well as complex viable tissue surrogates. A
finite-difference/front-tracking simulation model was
presented for deposition of viscous compound droplets onto
a receiving surface with the inclusion of significant
hydrodynamic pressures, capillary forces, and shear stresses.
Several parameters, such as Weber number, diameter
ratio, viscosity ratio, Reynolds number, surface tension ratio,
and equilibrium contact angle were investigated for their
influences on the transient deformation of a double
emulsion droplet during bio-printing. Such strategy can
accelerate the incorporation of 3D bioprinting technologies
into cancer research and the development of more precise
and reliable anticancer drug delivery systems.
Microfluidic channel combined with 3D printing is
emerging as a powerful tool in lab-on-a-chip and biological
study. With the fast technical development in the design
complexity, manufacture resolution and throughput, more
applications will bring the technology to wide acceptance
and commercialization. Overall, the development needs the
integration of multidisciplinary technologies in engineering,
biomaterials, cell biology, physics, and medicine [
This study was financially supported by Academic Research Fund (AcRF),
RG171/15, Ministry of Education, Singapore.
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1. Barry R , Ivanov D. Microfluidics in biotechnology . J Nanobiotechnol . 2004 ; 2 ( 1 ): 2 .
2. Dittrich PS , Manz A. Lab-on-a-chip: microfluidics in drug discovery . Nat Rev Drug Discov . 2006 ; 5 ( 3 ): 210 - 8 .
3. Guo L , Feng J , Fang Z , Xu J , Lu X . Application of microfluidic “lab-on-a-chip” for the detection of mycotoxins in foods . Trends Food Sci Technol . 2015 ; 46 ( 2 ): 252 - 63 .
4. Weibel DB , Whitesides GM . Applications of microfluidics in chemical biology . Curr Opin Chem Biol . 2006 ; 10 ( 6 ): 584 - 91 .
5. Whitesides GM . The origins and the future of microfluidics . Nature . 2006 ; 442 ( 7101 ): 368 - 73 .
6. Verpoorte E. Microfluidic chips for clinical and forensic analysis . Electrophoresis . 2002 ; 23 ( 5 ): 677 - 712 .
7. Benavente-Babace A , Gallego-Pérez D , Hansford DJ , Arana S , Pérez-Lorenzo E , Mujika M . Single-cell trapping and selective treatment via co-flow within a microfluidic platform . Biosens Bioelectron . 2014 ; 61 : 298 - 305 .
8. Holmes D , Whyte G , Bailey J , Vergara-Irigaray N , Ekpenyong A , Guck J , Duke T. Separation of blood cells with differing deformability using deterministic lateral displacement . Interface Focus . 2014 ; 4 ( 6 ): 20140011 .
9. Nge PN , Rogers CI , Woolley AT . Advances in microfluidic materials, functions, integration, and applications . Chem Rev . 2013 ; 113 ( 4 ): 2550 - 83 .
10. Moltzahn F , Olshen AB , Baehner L , Peek A , Fong L , Stöppler H , Simko J , Hilton JF , Carroll P , Blelloch R . Microfluidic-based multiplex qRT-PCR identifies diagnostic and prognostic microRNA signatures in the sera of prostate cancer patients . Cancer Res . 2011 ; 71 ( 2 ): 550 - 60 .
11. Kuo C-T , Chiang C-L , Chang C-H , Liu H-K , Huang G-S , Huang RY-J , Lee H , Huang C-S , Wo AM . Modeling of cancer metastasis and drug resistance via biomimetic nano-cilia and microfluidics . Biomaterials . 2014 ; 35 ( 5 ): 1562 - 71 .
12. Novik E , Maguire TJ , Chao P , Cheng K , Yarmush ML . A microfluidic hepatic coculture platform for cell-based drug metabolism studies . Biochem Pharmacol . 2010 ; 79 ( 7 ): 1036 - 44 .
13. Dishinger JF , Kennedy RT . Serial immunoassays in parallel on a microfluidic chip for monitoring hormone secretion from living cells . Anal Chem . 2007 ; 79 ( 3 ): 947 - 54 .
14. Liu P , Mathies RA . Integrated microfluidic systems for high-performance genetic analysis . Trends Biotechnol . 2009 ; 27 ( 10 ): 572 - 81 .
15. Haring A.P. , Sontheimer H. and Johnson B.N. Microphysiological human brain and neural systems-on-a-Chip: potential alternatives to small animal models and emerging platforms for drug discovery and personalized medicine . Stem cell reviews and reports:1-26 , 2017 .
16. Huh D , Hamilton GA , Ingber DE . From 3D cell culture to organs-on-chips . Trends Cell Biol . 2011 ; 21 ( 12 ): 745 - 54 .
17. Huh D , Matthews BD , Mammoto A , Montoya-Zavala M , Hsin HY , Ingber DE . Reconstituting organ-level lung functions on a chip . Science . 2010 ; 328 ( 5986 ): 1662 - 8 .
18. Kimura H , Yamamoto T , Sakai H , Sakai Y , Fujii T. An integrated microfluidic system for long-term perfusion culture and on-line monitoring of intestinal tissue models . Lab Chip . 2008 ; 8 ( 5 ): 741 - 6 .
19. Sackmann EK , Fulton AL , Beebe DJ . The present and future role of microfluidics in biomedical research . Nature . 2014 ; 507 ( 7491 ): 181 - 9 .
20. Tsai M , Kita A , Leach J , Rounsevell R , Huang JN , Moake J , Ware RE , Fletcher DA , Lam WA. In vitro modeling of the microvascular occlusion and thrombosis that occur in hematologic diseases using microfluidic technology . J Clin Invest . 2012 ; 122 ( 1 ): 408 - 18 .
21. Hanada Y , Sugioka K , Kawano H , Ishikawa IS , Miyawaki A , Midorikawa K. Nano-aquarium for dynamic observation of living cells fabricated by femtosecond laser direct writing of photostructurable glass . Biomed Microdevices . 2008 ; 10 ( 3 ): 403 - 10 .
22. Hanada Y , Sugioka K , Shihira-Ishikawa I , Kawano H , Miyawaki A , Midorikawa K. 3D microfluidic chips with integrated functional microelements fabricated by a femtosecond laser for studying the gliding mechanism of cyanobacteria . Lab Chip . 2011 ; 11 ( 12 ): 2109 - 15 .
23. Duffy DC , McDonald JC , Schueller OJ , Whitesides GM . Rapid prototyping of microfluidic systems in poly (dimethylsiloxane) . Anal Chem . 1998 ; 70 ( 23 ): 4974 - 84 .
24. Amin R , Knowlton S , Hart A , Yenilmez B , Ghaderinezhad F , Katebifar S , Messina M , Khademhosseini A , Tasoglu S. 3D-printed microfluidic devices . Biofabrication . 2016 ; 8 ( 2 ): 022001 .
25. Au A.K. , Huynh W. , Horowitz L.F. and Folch A. 3D-printed microfluidics . Angewandte Chemie International Edition , 2016 .
26. Tseng P , Murray C , Kim D , Di Carlo D. Research highlights: printing the future of microfabrication . Lab Chip . 2014 ; 14 ( 9 ): 1491 - 5 .
27. Rogers CI , Qaderi K , Woolley AT , Nordin GP . 3D printed microfluidic devices with integrated valves . Biomicrofluidics . 2015 ; 9 ( 1 ): 016501 .
28. Fuard D , Tzvetkova-Chevolleau T , Decossas S , Tracqui P , Schiavone P . Optimization of poly-di-methyl-siloxane (PDMS) substrates for studying cellular adhesion and motility . Microelectron Eng . 2008 ; 85 ( 5 ): 1289 - 93 .
29. McDonald JC , Whitesides GM . Poly (dimethylsiloxane) as a material for fabricating microfluidic devices . Acc Chem Res . 2002 ; 35 ( 7 ): 491 - 9 .
30. Zhang Q , Austin RH . Applications of microfluidics in stem cell biology . BioNanoScience . 2012 ; 2 ( 4 ): 277 - 86 .
31. Folch A . Introduction to bioMEMS: CRC Press; 2016 .
32. Au AK , Huynh W , Horowitz LF , Folch A . 3D-printed microfluidics . Angew Chem Int Ed . 2016 ; 55 ( 12 ): 3862 - 81 .
33. Macdonald NP , Zhu F , Hall C , Reboud J , Crosier P , Patton E , Wlodkowic D , Cooper J . Assessment of biocompatibility of 3D printed photopolymers using zebrafish embryo toxicity assays . Lab Chip . 2016 ; 16 ( 2 ): 291 - 7 .
34. Lee KY , Mooney DJ . Hydrogels for tissue engineering . Chem Rev . 2001 ; 101 ( 7 ): 1869 - 80 .
35. Murphy SV , Atala A. 3D bioprinting of tissues and organs . Nat Biotechnol . 2014 ; 32 ( 8 ): 773 - 85 .
36. Ling Y , Rubin J , Deng Y , Huang C , Demirci U , Karp JM , Khademhosseini A. A cell-laden microfluidic hydrogel . Lab Chip . 2007 ; 7 ( 6 ): 756 - 62 .
37. Bertassoni LE , Cecconi M , Manoharan V , Nikkhah M , Hjortnaes J , Cristino AL , Barabaschi G , Demarchi D , Dokmeci MR , Yang Y . Hydrogel bioprinted microchannel networks for vascularization of tissue engineering constructs . Lab Chip . 2014 ; 14 ( 13 ): 2202 - 11 .
38. Stanton M , Samitier J , Sánchez S. Bioprinting of 3D hydrogels . Lab Chip . 2015 ; 15 ( 15 ): 3111 - 5 .
39. Au AK , Lee W , Folch A . Mail-order microfluidics: evaluation of stereolithography for the production of microfluidic devices . Lab Chip . 2014 ; 14 ( 7 ): 1294 - 301 .
40. Gowers SA , Curto VF , Seneci CA , Wang C , Anastasova S , Vadgama P , Yang G-Z , Boutelle MG. 3D printed microfluidic device with integrated biosensors for online analysis of subcutaneous human microdialysate . Anal Chem . 2015 ; 87 ( 15 ): 7763 - 70 .
41. Comina G , Suska A , Filippini D. PDMS lab-on-a-chip fabrication using 3D printed templates . Lab Chip . 2014 ; 14 ( 2 ): 424 - 30 .
42. Lee JM , Zhang M , Yeong WY . Characterization and evaluation of 3D printed microfluidic chip for cell processing . Microfluid Nanofluid . 2016 ; 20 ( 1 ): 5 .
43. Hwang Y , Paydar OH , Candler RN . 3D printed molds for non-planar PDMS microfluidic channels . Sensors Actuators A Phys . 2015 ; 226 : 137 - 42 .
44. Song S-H , Lee C-K , Kim T-J , Shin I-C , Jun S-C, Jung H-I. A rapid and simple fabrication method for 3-dimensional circular microfluidic channel using metal wire removal process . Microfluid Nanofluid . 2010 ; 9 ( 2 -3): 533 - 40 .
45. Verma MK , Majumder A , Ghatak A . Embedded template-assisted fabrication of complex microchannels in PDMS and design of a microfluidic adhesive . Langmuir . 2006 ; 22 ( 24 ): 10291 - 5 .
46. Bellan LM , Singh SP , Henderson PW , Porri TJ , Craighead HG , Spector JA . Fabrication of an artificial 3-dimensional vascular network using sacrificial sugar structures . Soft Matter . 2009 ; 5 ( 7 ): 1354 - 7 .
47. Lee J , Paek J , Kim J . Sucrose-based fabrication of 3D-networked, cylindrical microfluidic channels for rapid prototyping of lab-on-a-chip and vasomimetic devices . Lab Chip . 2012 ; 12 ( 15 ): 2638 - 42 .
48. Miller JS , Stevens KR , Yang MT , Baker BM , Nguyen D-HT , Cohen DM , Toro E , Chen AA , Galie PA , Yu X . Rapid casting of patterned vascular networks for perfusable engineered three-dimensional tissues . Nat Mater . 2012 ; 11 ( 9 ): 768 - 74 .
49. Yang J , Li K , Zhu L , Tang W. Fabrication of PDMS microfluidic devices with 3D wax jetting . RSC Adv . 2017 ; 7 ( 6 ): 3313 - 20 .
50. Hansen CJ , Wu W , Toohey KS , Sottos NR , White SR , Lewis JA . Self-healing materials with interpenetrating microvascular networks . Adv Mater . 2009 ; 21 ( 41 ): 4143 - 7 .
51. Therriault D , White SR , Lewis JA . Chaotic mixing in three-dimensional microvascular networks fabricated by direct-write assembly . Nat Mater . 2003 ; 2 ( 4 ): 265 - 71 .
52. Wu W , Hansen CJ , Aragón AM , Geubelle PH , White SR , Lewis JA . Direct-write assembly of biomimetic microvascular networks for efficient fluid transport . Soft Matter . 2010 ; 6 ( 4 ): 739 - 42 .
53. Wu W , DeConinck A , Lewis JA . Omnidirectional printing of 3D microvascular networks . Adv Mater . 2011 ; 23 ( 24 ): H178 - 83 .
54. Schwartz MA , Chen CS . Deconstructing dimensionality . Science . 2013 ; 339 ( 6118 ): 402 - 4 .
55. Boland T , Xu T , Damon B , Cui X . Application of inkjet printing to tissue engineering . Biotechnol J. 2006 ; 1 ( 9 ): 910 - 7 .
56. Nakamura M , Kobayashi A , Takagi F , Watanabe A , Hiruma Y , Ohuchi K , Iwasaki Y , Horie M , Morita I , Takatani S. Biocompatible inkjet printing technique for designed seeding of individual living cells . Tissue Eng . 2005 ; 11 ( 11 -12): 1658 - 66 .
57. Zein I , Hutmacher DW , Tan KC , Teoh SH . Fused deposition modeling of novel scaffold architectures for tissue engineering applications . Biomaterials . 2002 ; 23 ( 4 ): 1169 - 85 .
58. Chisti Y. Hydrodynamic damage to animal cells . Crit Rev Biotechnol . 2001 ; 21 ( 2 ): 67 - 110 .
59. Sechi D , Greer B , Johnson J , Hashemi N. Three-dimensional paper-based microfluidic device for assays of protein and glucose in urine . Anal Chem . 2013 ; 85 ( 22 ): 10733 - 7 .
60. Thomas MS , Millare B , Clift JM , Bao D , Hong C , Vullev VI . Print-and-peel fabrication for microfluidics: what's in it for biomedical applications ? Ann Biomed Eng. 2010 ; 38 ( 1 ): 21 - 32 .
61. Erkal JL , Selimovic A , Gross BC , Lockwood SY , Walton EL , McNamara S , Martin RS , Spence DM . 3D printed microfluidic devices with integrated versatile and reusable electrodes . Lab Chip . 2014 ; 14 ( 12 ): 2023 - 32 .
62. Kitson PJ , Rosnes MH , Sans V , Dragone V , Cronin L. Configurable 3D -printed millifluidic and microfluidic 'lab on a chip'reactionware devices . Lab Chip . 2012 ; 12 ( 18 ): 3267 - 71 .
63. Chiu DT , Jeon NL , Huang S , Kane RS , Wargo CJ , Choi IS , Ingber DE , Whitesides GM . Patterned deposition of cells and proteins onto surfaces by using three-dimensional microfluidic systems . Proc Natl Acad Sci . 2000 ; 97 ( 6 ): 2408 - 13 .
64. El-Ali J , Sorger PK , Jensen KF . Cells on chips . Nature . 2006 ; 442 ( 7101 ): 403 - 11 .
65. Kamei K. Cutting-edge microfabricated biomedical tools for human pluripotent stem cell research . J Lab Automation . 2013 ; 18 ( 6 ): 469 - 81 .
66. Kim D-H , Beebe DJ , Levchenko A . Micro-and nanoengineering for stem cell biology: the promise with a caution . Trends Biotechnol . 2011 ; 29 ( 8 ): 399 - 408 .
67. Kamei K , Mashimo Y , Koyama Y , Fockenberg C , Nakashima M , Nakajima M , Li J , Chen Y. 3D printing of soft lithography mold for rapid production of polydimethylsiloxane-based microfluidic devices for cell stimulation with concentration gradients . Biomed Microdevices . 2015 ; 17 ( 2 ): 36 .
68. Connell JL , Ritschdorff ET , Whiteley M , Shear JB . 3D printing of microscopic bacterial communities . Proc Natl Acad Sci . 2013 ; 110 ( 46 ): 18380 - 5 .
69. Leonard P , Hearty S , Brennan J , Dunne L , Quinn J , Chakraborty T , O'Kennedy R. Advances in biosensors for detection of pathogens in food and water . Enzym Microb Technol . 2003 ; 32 ( 1 ): 3 - 13 .
70. Lee W , Kwon D , Choi W , Jung GY , Au AK , Folch A , Jeon S. 3D-printed microfluidic device for the detection of pathogenic bacteria using size-based separation in helical channel with trapezoid cross-section . Sci Rep . 2015 ; 5 : 7717 .
71. White SR , Sottos N , Geubelle P , Moore J , Kessler MR , Sriram S , Brown E , Viswanathan S. Autonomic healing of polymer composites . Nature . 2001 ; 409 ( 6822 ): 794 - 7 .
72. Borenstein JT , Weinberg EJ , Orrick BK , Sundback C , Kaazempur-Mofrad MR , Vacanti JP . Microfabrication of three-dimensional engineered scaffolds . Tissue Eng . 2007 ; 13 ( 8 ): 1837 - 44 .
73. Norotte C , Marga FS , Niklason LE , Forgacs G . Scaffold-free vascular tissue engineering using bioprinting . Biomaterials . 2009 ; 30 ( 30 ): 5910 - 7 .
74. Williamson A , Singh S , Fernekorn U , Schober A . The future of the patientspecific body-on-a-chip . Lab Chip . 2013 ; 13 ( 18 ): 3471 - 80 .
75. Wheeler TD , Stroock AD . The transpiration of water at negative pressures in a synthetic tree . Nature . 2008 ; 455 ( 7210 ): 208 - 12 .
76. Toohey KS , Hansen CJ , Lewis JA , White SR , Sottos NR . Delivery of two-part self-healing chemistry via microvascular networks . Adv Funct Mater . 2009 ; 19 ( 9 ): 1399 - 405 .
77. Kaully T , Kaufman-Francis K , Lesman A , Levenberg S . Vascularization-the conduit to viable engineered tissues . Tissue Eng B Rev . 2009 ; 15 ( 2 ): 159 - 69 .
78. Du Y , Lo E , Ali S , Khademhosseini A . Directed assembly of cell-laden microgels for fabrication of 3D tissue constructs . Proc Natl Acad Sci . 2008 ; 105 ( 28 ): 9522 - 7 .
79. Khademhosseini A , Langer R , Borenstein J , Vacanti JP . Microscale technologies for tissue engineering and biology . Proc Natl Acad Sci U S A . 2006 ; 103 ( 8 ): 2480 - 7 .
80. Paul A , Hasan A , Rodes L , Sangaralingam M , Prakash S . Bioengineered baculoviruses as new class of therapeutics using micro and nanotechnologies: principles, prospects and challenges . Adv Drug Deliv Rev . 2014 ; 71 : 115 - 30 .
81. Choi NW , Cabodi M , Held B , Gleghorn JP , Bonassar LJ , Stroock AD . Microfluidic scaffolds for tissue engineering . Nat Mater . 2007 ; 6 ( 11 ): 908 - 15 .
82. Park JH , Chung BG , Lee WG , Kim J , Brigham MD , Shim J , Lee S , Hwang CM , Durmus NG , Demirci U . Microporous cell-laden hydrogels for engineered tissue constructs . Biotechnol Bioeng . 2010 ; 106 ( 1 ): 138 - 48 .
83. Bae H , Puranik AS , Gauvin R , Edalat F , Carrillo-Conde B , Peppas NA , Khademhosseini A . Building vascular networks . Sci Transl Med . 2012 ; 4 ( 160 ): 160ps123 - 3 .
84. Folkman J. Angiogenesis in cancer, vascular, rheumatoid and other disease . Nat Med . 1995 ; 1 ( 1 ): 27 - 30 .
85. Lee VK , Kim DY , Ngo H , Lee Y , Seo L , Yoo S-S , Vincent PA , Dai G . Creating perfused functional vascular channels using 3D bio-printing technology . Biomaterials . 2014 ; 35 ( 28 ): 8092 - 102 .
86. Attalla R , Ling C , Selvaganapathy P. Fabrication and characterization of gels with integrated channels using 3D printing with microfluidic nozzle for tissue engineering applications . Biomed Microdevices . 2016 ; 18 ( 1 ): 1 - 12 .
87. Hong S , Sycks D , Chan HF , Lin S , Lopez GP , Guilak F , Leong KW , Zhao X . 3D printing of highly stretchable and tough hydrogels into complex, cellularized structures . Adv Mater . 2015 ; 27 ( 27 ): 4035 - 40 .
88. Colosi C , Shin SR , Manoharan V , Massa S , Costantini M , Barbetta A , Dokmeci MR , Dentini M , Khademhosseini A . Microfluidic bioprinting of heterogeneous 3D tissue constructs using low-viscosity bioink . Adv Mater . 2016 ; 28 ( 4 ): 677 - 84 .
89. Knowlton S , Joshi A , Yenilmez B , Ozbolat IT , Chua CK , Khademhosseini A , Tasoglu S. Advancing cancer research using bioprinting for tumor-on-a-chip platforms . Int J Bioprinting . 2016 ; 2 ( 2 )
90. Zhao Y , Yao R , Ouyang L , Ding H , Zhang T , Zhang K , Cheng S , Sun W. Three-dimensional printing of Hela cells for cervical tumor model in vitro . Biofabrication . 2014 ; 6 ( 3 ): 035001 .
91. Knowlton S , Onal S , Yu CH , Zhao JJ , Tasoglu S . Bioprinting for cancer research . Trends Biotechnol . 2015 ; 33 ( 9 ): 504 - 13 .
92. Bhise NS , Manoharan V , Massa S , Tamayol A , Ghaderi M , Miscuglio M , Lang Q , Zhang YS , Shin SR , Calzone G. A liver-on-a-chip platform with bioprinted hepatic spheroids . Biofabrication . 2016 ; 8 ( 1 ): 014101 .
93. Knowlton S , Yenilmez B , Tasoglu S . Towards single-step biofabrication of organs on a chip via 3D printing . Trends Biotechnol . 2016 ; 34 ( 9 ): 685 - 8 .
94. Singh M , Tong Y , Webster K , Cesewski E , Haring AP , Laheri S , Carswell B , O'Brien TJ , Aardema CH , Senger RS . 3D printed conformal microfluidics for isolation and profiling of biomarkers from whole organs . Lab Chip . 2017 ; 17 ( 15 ): 2561 - 71 .
95. Illg T , Löb P , Hessel V . Flow chemistry using milli-and microstructured reactors-from conventional to novel process windows . Bioorg Med Chem . 2010 ; 18 ( 11 ): 3707 - 19 .
96. Kitson PJ , Glatzel S , Chen W , Lin C-G , Song Y-F , Cronin L. 3D printing of versatile reactionware for chemical synthesis . Nat Protoc . 2016 ; 11 ( 5 ): 920 - 36 .
97. Okafor O. , Weilhard A. , Fernandes J.A. , Karjalainen E. , Goodridge R. and Sans V. Advanced reactor engineering with 3D printing for the continuous-flow synthesis of silver nanoparticles . Reaction Chemistry & Engineering , 2017 .
98. Ho CMB , Ng SH , Li KHH , Yoon Y -J. 3D printed microfluidics for biological applications . Lab Chip . 2015 ; 15 ( 18 ): 3627 - 37 .
99. Zavorotnitsienko D. Understanding 3D printer quality & resolution . http:// ilios3d.com/en/product-documentation/ilios-documentation - 3dprint-quality
100. Bhargava KC , Thompson B , Malmstadt N. Discrete elements for 3D microfluidics . Proc Natl Acad Sci . 2014 ; 111 ( 42 ): 15013 - 8 .
101. Tumbleston JR , Shirvanyants D , Ermoshkin N , Janusziewicz R , Johnson AR , Kelly D , Chen K , Pinschmidt R , Rolland JP , Ermoshkin A . Continuous liquid interface production of 3D objects . Science . 2015 ; 347 ( 6228 ): 1349 - 52 .
102. Yazdi AA , Popma A , Wong W , Nguyen T , Pan Y , Xu J. 3D printing: an emerging tool for novel microfluidics and lab-on-a-chip applications . Microfluid Nanofluid . 2016 ; 20 ( 3 ): 1 - 18 .
103. Tsuda S , Jaffery H , Doran D , Hezwani M , Robbins PJ , Yoshida M , Cronin L . Customizable 3D printed 'plug and play'millifluidic devices for programmable fluidics . PLoS One . 2015 ; 10 ( 11 ): e0141640 .
104. Meng Z-J , Wang W , Liang X , Zheng W-C , Deng N-N , Xie R , Ju X-J , Liu Z , Chu L-Y. Plug -n-play microfluidic systems from flexible assembly of glass-based flow-control modules . Lab Chip . 2015 ; 15 ( 8 ): 1869 - 78 .
105. Gong H , Woolley AT , Nordin GP . High density 3D printed microfluidic valves, pumps, and multiplexers . Lab Chip . 2016 ; 16 ( 13 ): 2450 - 8 .
106. Waheed S , Cabot JM , Macdonald NP , Lewis T , Guijt RM , Paull B , Breadmore MC . 3D printed microfluidic devices: enablers and barriers . Lab Chip . 2016 ; 16 ( 11 ): 1993 - 2013 .
107. Bhattacharjee N , Urrios A , Kang S , Folch A . The upcoming 3D-printing revolution in microfluidics . Lab Chip . 2016 ; 16 ( 10 ): 1720 - 42 .
108. O 'Neill P , Ben AA , Vázquez M , Liu J , Marczak S , Slouka Z , Chang HC , Diamond D , Brabazon D. Advances in three-dimensional rapid prototyping of microfluidic devices for biological applications . Biomicrofluidics . 2014 ; 8 ( 5 ): 052112 .
109. Li Y-SJ , Haga JH , Chien S. Molecular basis of the effects of shear stress on vascular endothelial cells . J Biomech . 2005 ; 38 ( 10 ): 1949 - 71 .
110. Yan Y , Wang X , Pan Y , Liu H , Cheng J , Xiong Z , Lin F , Wu R , Zhang R , Lu Q. Fabrication of viable tissue-engineered constructs with 3D cell-assembly technique . Biomaterials . 2005 ; 26 ( 29 ): 5864 - 71 .
111. Kim JB . Three-dimensional tissue culture models in cancer biology . Semin Cancer Biol. Elsevier . 2005 : 365 - 77 .
112. Vargo-Gogola T , Rosen JM . Modelling breast cancer: one size does not fit all . Nat Rev Cancer . 2007 ; 7 ( 9 ): 659 - 72 .
113. Chan CY , Huang P-H , Guo F , Ding X , Kapur V , Mai JD , Yuen PK , Huang TJ . Accelerating drug discovery via organs-on-chips . Lab Chip . 2013 ; 13 ( 24 ): 4697 - 710 .
114. Chang R , Emami K , Wu H , Sun W. Biofabrication of a three-dimensional liver micro-organ as an in vitro drug metabolism model . Biofabrication . 2010 ; 2 ( 4 ): 045004 .
115. Yao R , Zhang R , Yan Y , Wang X. In vitro angiogenesis of 3D tissue engineered adipose tissue . J Bioact Compat Polym . 2009 ; 24 ( 1 ): 5 - 24 .
116. Fedorovich NE , De Wijn JR , Verbout AJ , Alblas J , Dhert WJ . Threedimensional fiber deposition of cell-laden, viable, patterned constructs for bone tissue printing . Tissue Eng A . 2008 ; 14 ( 1 ): 127 - 33 .
117. Li S , Xiong Z , Wang X , Yan Y , Liu H , Zhang R . Direct fabrication of a hybrid cell/hydrogel construct by a double-nozzle assembling technology . J Bioact Compat Polym . 2009 ; 24 ( 3 ): 249 - 65 .
118. Knowlton S , Yu CH , Ersoy F , Emadi S , Khademhosseini A , Tasoglu S. 3Dprinted microfluidic chips with patterned, cell-laden hydrogel constructs . Biofabrication . 2016 ; 8 ( 2 ): 025019 .
119. Hamid Q , Wang C , Snyder J , Williams S , Liu Y , Sun W. Maskless fabrication of cell-laden microfluidic chips with localized surface functionalization for the co-culture of cancer cells . Biofabrication . 2015 ; 7 ( 1 ): 015012 .
120. Hansen CJ , Saksena R , Kolesky DB , Vericella JJ , Kranz SJ , Muldowney GP , Christensen KT , Lewis JA . High-throughput printing via microvascular multinozzle arrays . Adv Mater . 2013 ; 25 ( 1 ): 96 - 102 .
121. Kolesky DB , Truby RL , Gladman A , Busbee TA , Homan KA , Lewis JA. 3D bioprinting of vascularized, heterogeneous cell-laden tissue constructs . Adv Mater . 2014 ; 26 ( 19 ): 3124 - 30 .
122. Lind JU , Busbee TA , Valentine AD , Pasqualini FS , Yuan H , Yadid M , Park S-J , Kotikian A , Nesmith AP , Campbell PH . Instrumented cardiac microphysiological devices via multimaterial three-dimensional printing . Nat Mater . 2017 ; 16 ( 3 ): 303 - 8 .
123. Tasoglu S , Demirci U . Bioprinting for stem cell research . Trends Biotechnol . 2013 ; 31 ( 1 ): 10 - 9 .
124. Chan HN , Chen Y , Shu Y , Chen Y , Tian Q , Wu H . Direct, one-step molding of 3D-printed structures for convenient fabrication of truly 3D PDMS microfluidic chips . Microfluid Nanofluid . 2015 ; 19 ( 1 ): 9 - 18 .
125. He Y , Qiu J , Fu J , Zhang J , Ren Y , Liu A . Printing 3D microfluidic chips with a 3D sugar printer . Microfluid Nanofluid . 2015 ; 19 ( 2 ): 447 - 56 .
126. Parekh DP , Ladd C , Panich L , Moussa K , Dickey MD . 3D printing of liquid metals as fugitive inks for fabrication of 3D microfluidic channels . Lab Chip . 2016 ; 16 ( 10 ): 1812 - 20 .
127. Johnson BN , Lancaster KZ , Hogue IB , Meng F , Kong YL , Enquist LW , McAlpine MC . 3D printed nervous system on a chip . Lab Chip . 2016 ; 16 ( 8 ): 1393 - 400 .
128. Lu Y , Shi W , Qin J , Lin B. Fabrication and characterization of paper-based microfluidics prepared in nitrocellulose membrane by wax printing . Anal Chem . 2009 ; 82 ( 1 ): 329 - 35 .
129. Shen W , Li M , Ye C , Jiang L , Song Y . Direct-writing colloidal photonic crystal microfluidic chips by inkjet printing for label-free protein detection . Lab Chip . 2012 ; 12 ( 17 ): 3089 - 95 .
130. Park C , Lee J , Kim Y , Kim J , Lee J , Park S. 3D-printed microfluidic magnetic preconcentrator for the detection of bacterial pathogen using an ATP luminometer and antibody-conjugated magnetic nanoparticles . J Microbiol Methods . 2017 ; 132 : 128 - 33 .
131. Zhang B , Montgomery M , Chamberlain MD , Ogawa S , Korolj A , Pahnke A , Wells LA , Massé S , Kim J , Reis L . Biodegradable scaffold with built-in vasculature for organ-on-a-chip engineering and direct surgical anastomosis . Nat Mater . 2016 ; 15 ( 6 ): 669 .