This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
1
Bioprinting towards Organ Fabrication:
Challenges and Future Trends
Ibrahim Ozbolat and Yin Yu

Abstract— Tissue engineering has been a promising field of
research, offering hope for bridging the gap between organ
shortage and transplantation needs. However, building threedimensional (3D) vascularized organs remains the main
technological barrier to be overcome. Organ printing, which is
defined as computer-aided additive biofabrication of 3D cellular
tissue constructs, has shed light on advancing this field into a new
era. Organ printing takes advantage of rapid prototyping (RP)
technology to print cells, biomaterials, and cell-laden biomaterials
individually or in tandem, layer by layer, directly creating 3D
tissue-like structures. Here, we overview RP-based bioprinting
approaches and discuss the current challenges and trends towards
fabricating living organs for transplant in the near future.
proliferation, and direct cell differentiation toward certain
lineages. Although significant success has been achieved in the
past decades both in research and clinical applications [7], it is
obvious that complex 3D organs require more precise multicellular structures with vascular network integration, which
cannot be fulfilled by traditional methods.
Index Terms— Bioadditive manufacturing, bioprinting, organ
fabrication, tissue engineering
I. INTRODUCTION
O
shortage has become more problematic in spite of
an increase in willing donors. From July 2000 to July
2001, for example, approximately 80,000 people in the United
States awaited an organ transplant, with less than a third
receiving it [1]. The solution to this problem, as with the
solutions to other grand engineering challenges, requires longterm solutions by building or manufacturing living organs from
a person’s own cells. For the past three decades, tissue
engineering has emerged as a multidisciplinary field involving
scientists, engineers, and physicians, for the purpose of
creating biological substitutes mimicking native tissue to
replace damaged tissues or restore malfunctioning organs [2].
The traditional tissue engineering strategy is to seed cells onto
scaffolds, which can then direct cell proliferation and
differentiation into three-dimensional (3D) functioning tissues.
Both synthetic and natural polymers have been used to
engineer various tissue grafts like skin, cartilage, bone, and
bladder [3-6]. To be successfully used for tissue engineering,
these materials must be biocompatible and biodegradable, with
the mechanical strength to support cell attachment,
RGAN
Manuscript received December 12, 2012. This research was supported by
National Institutes of Health (NIH) and the Institute for Clinical and
Translational Science (ICTS) grant number ULIRR024979.
I.T. Ozbolat is with The University of Iowa Mechanical and Industrial
Engineering and Biomanufacturing Laboratory, Iowa City, IA, 52242 USA
(corresponding author to provide phone: 319-384-0810; fax: 319-335-5669;
e-mail: [email protected]).
Y. Yin is with The University of Iowa Biomedical Engineering
Department and Biomanufacturing Laboratory, Iowa City, IA 52242 USA (email: yin-yu@ uiowa.edu).
Fig. 1: Organ printing concept: 3D functional organs are printed from
the bottom up using living cells in a supportive medium (image
courtesy of Christopher Barnatt [8]).
A computer-aided bioadditive manufacturing process has
emerged to deposit living cells together with hydrogel-base
scaffolds for 3D tissue and organ fabrication. Bioprinting or
direct cell printing is an extension of tissue engineering, as it
intends to create de novo organs. It uses bioadditive
manufacturing technologies, including laser-based writing [9],
inkjet-based printing [10] and extrusion-based deposition [11].
Bioprinting offers great precision on spatial placement of the
cells themselves, rather than providing scaffold support alone
[12]. Although still in its infancy, this technology appears to be
more promising for advancing tissue engineering towards
organ fabrication, ultimately mitigating organ shortage and
saving lives. Figure 1 demonstrates the concept of futuristic
3D direct organ printing technology, where multiple living
cells with the supportive media stored in cartridges are printed
layer by layer using inkjet printing technology. It offers a
controllable fabrication process, which allows precise
placement of various biomaterial and/or cell types
simultaneously according to the natural compartments of the
target tissue or organs. Multiple cell types, including organspecific cells and blood vessel cells, i.e., smooth muscle and
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
endothelial cells (ECs), constitute the entire organ. Although
the concept seems to be trivial considering the complexity and
functionality of the parts that can be manufactured using
contemporary rapid prototyping (RP) technology [13], several
challenges impede the evolution of organ printing. This paper
discusses the current state of the art in bioprinting technology,
its advantages and limitations, and grand challenges and future
trends towards fabrication of living organs in both research
and clinical scenarios.
II. BIOPRINTING: CURRENT STATE OF THE ART
Bioprinting, where living cells are precisely printed in a
certain pattern, has great potential and promise for fabricating
engineered living organs. Based on their working principles,
bioprinting systems can be primarily classified as: (1) laserbased, (2) inkjet-based, or (3) extrusion-based.
Fig. 2: Bioprinting techniques including (A) laser-based writing of
cells, (B) ink-jet based systems and (C) extrusion-based deposition.
Laser technology has recently been applied in the cell
printing process, in which laser energy is used to excite the
cells and give patterns to control spatially the cellular
environment. A laser-based system was first introduced in
1999 by Odde et al. to process 2D cell patterning [14]. Laser
direct-write (LDW) is a biofabrication method capable of
rapidly creating precise patterns of viable cells on petri dishes.
In LDW, cells suspended in a solution in donor slides are
transferred to a collector slide using laser energy. A laser pulse
creates a bubble, and shock waves are generated by the bubble
formation, which eventually propel cells towards the collector
substrate (see Fig. 2(A)). Micro-scale cell patterning can be
achieved through optimizing viscosity of biological material
(bioink), laser printing speed, laser energy, and pulse
frequency [15]. Writing of multiple cell types is also feasible
by selectively propelling different cells to the collector
substrate. Laser printing technology is also integrated with
scaffold printing, where LDW is performed in tandem with
photopolymerization of hydrogels. It basically deposits cells in
2
a certain pattern onto a substrate by a laser beam. This is
followed by deposition of hydrogel on top of each layer of
cells, and the process is repeated for multiple cycles to get a
3D structure. Nahmias et al. successfully performed
hepatocytes patterning in collagen and MatrigelTM using laserguided 3D cell writing [16]. In their study, three layers of cells
and hydrogels were alternately deposited on top of each other,
forming a 3D cellular structure. Cell viability and proliferation
was well-maintained post-deposition.
Inkjet-based bioprinting was introduced in the early 2000s
and built a great foundation for future organ printing
technologies. In this technique, living cells are printed in the
form of droplets through cartridges instead of seeding them on
scaffolds (see Fig. 2(B)). It uses a non-contact reprographic
technique that takes digital data from a computer representing
tissue or organs, and reproduces it onto a substrate using “bioink” made of cells and biomaterials [10]. Boland et al. used a
thermal inkjet printer to successfully fabricate 3D cellular
assemblies of bovine aortal ECs with thermosensitive gels
[10]. Post-incubation, printed structures showed high cell
viability and maintained cell phenotype. Cui and his coworkers
[17] applied inkjet printing technology to repair human
articular cartilage, showing its promising potential for highefficiency direct tissue regeneration. Huang and his coworkers
[18] developed a bipolar wave-based drop-on-demand jetting.
In their studies, cell-encapsulated alginate microspheres were
jetted and assembled to create vertically oriented, short,
tubular structures [19]. Inkjet-based system allows printing
single cells or cell aggregates [20, 21] by controlling process
parameters such as cell concentration, drop volume, resolution,
nozzle diameter and average diameter of printed cells [22].
Weiss and his coworkers [23] developed a multi-head inkjetbased bioprinting platform for fabricating heterogeneous
structures with a concentration gradient changing from the
bottom up. Multiple growth factors such as fibrinogen and
thrombin and cells were printed with spatial precision in a
functionally graded manner into rat calvarial defect in-situ
[24]. They demonstrated the feasibility of in-situ printing;
however, this technique did not seem to be a practical
approach for clinicians due to complex nature of the process in
their study [24].
Another bioprinting technique has been introduced for
printing living cells and is based on the extrusion of
continuous filaments made of biomaterials. It is a combination
of a fluid-dispensing system and an automated three-axis
robotic system for extrusion and printing, respectively [11].
During printing, biomaterial is dispensed by a pressureassisted system, under the control of “robots,” resulting in
precise deposition of cells encapsulated in the cylindrical
filaments of desired 3D structures (see Fig. 2(C)). Wang et al.
used a 3D syringe-based bioprinting system to deposit
different cells with various biocompatible hydrogels [25, 26].
They used hepatocytes and adipose-derived stromal cells
(ADSCs) together with gelatin/chitosan hydrogels to engineer
an artificial liver. Sun and his coworkers [27-29] built a multi-
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
nozzle bioprinting system with the capacity to simultaneously
deposit cells and multiple biomaterials. Their rheology study
and cell viability assay were performed to investigate
mechanical-stress-induced cell damage during the printing
process [30]. The results showed that cell viability was
influenced by material flow rate, material concentration,
dispensing pressure, and nozzle geometry. Their findings can
serve as a guideline for future studies and optimization of the
deposition system. Kachouie et al. proposed a method using
hydrogel-encapsulated cells as tissue units to make a construct
with geometric patterns specific to target tissue types [31].
Although bioprinting is a promising way as a methodical
interface between tissue and engineering, each technique has
its own limitations. Laser-based systems have high resolution
and enable precise patterning of living cells, where cells can be
maintained registry within 5.6±2.5µm to the initial pattern
[32]. This resolution, certainly, cannot be achieved by other
bioprinting techniques that makes laser-based cell writing as a
great potential for micro-cellular features in organs and tissues
i.e. micro-vascularate. In the authors opinion, prolonged
fabrication time, laser shock related thermally and
mechanically induced cell deformation, interactions of cell
components with light, gravitational and random setting of
cells in the precursor solution, limitations in printing in third
dimension and the need for photo-crosslinkable biomaterials
should be overcome for future developments in laser-based
systems [15, 33, 34]. In order to improve resolution further
and increase throughput, parameters related to laser pulse
characteristics (i.e. pulse duration, wavelength, repetition rate,
energy and beam focus diameter), precursor solution
properties such as viscosity, thickness and surface tension and
substrate properties should be optimized [34]. One-directional
propulsion of cells toward the collector substrate (top to
bottom) certainly limits development of 3D structures with
complex heterocellular architectures. In order to expand this
technology in the third dimension, a rotating donor-side
carousel leveling system with rotational and linear stages can
be developed that allows printing multiple cells in different
layers for the development of heterocellular structures. Similar
technology has been recently demonstrated with multi-material
stereolithography process [35]. In addition, the fabrication
speed can be improved by increasing the laser pulse rate or
integrating multiple laser beams [36].
Inkjet-based systems are versatile and affordable that favors
cells encapsulation. For instance, Xu et al. [37] developed a
fabrication platform with a different working principle as
opposed to traditional inkjet printers, where droplets of
crosslinker chemical were deposited onto a suspension of
cardiomyocytes and alginate for 3D cardiac pseudo tissue
fabrication. In addition, the traditional setup allows integrating
multiple print heads easily to deposit multiple cell types, which
is one of the pivotal steps in fabricating heterocellular tissues
and organs. The other advantage of inkjet printing is that
surfaces where the cells are printed and patterned do not have
to be flat that favors cell printing in situ [38]. In general, all
3
other techniques require gentle handling of a flat deposition
surface that limits their direct applicability in surgery rooms.
Despite their great advantages, inkjet printers suffer from
drawbacks including significant cell damage and death as well
as cell sedimentation and aggregation due to small orifice
diameter that restricts printing cells in high densities (<5x106
cell/ml) [22, 39, 40]. In addition, structural integrity of the
printed structures is another obstacle, where the droplets do
not fuse into each other easily and the shape of droplets cannot
be controlled precisely. Adequate structural integrity is crucial
to retain the designed and the fabricated shape.
Extrusion-based systems provide relatively better structural
integrity due to continuous deposition of cylindrical struts. It is
the most convenient technique in rapidly fabricating 3D
porous cellular structures [12]. Although this technology lays
foundation for cell patterning for scale-up tissue and organ
fabrication technologies, it constitutes several limitations such
as shear stress induced cell deformation and limited material
selection due to need for rapid encapsulation of cells via
gelation. Shear stress on nozzle tip wall induces significant
drop in the number of living cells when the cell density is high;
however, this can be partially alleviated using optimum
process parameters such as biomaterial concentration, nozzle
pressure (ideally minimum), nozzle diameter and loaded cell
density [41]. Restricted biomaterial selection and low
resolution and accuracy brings limited applicability of
extrusion-based systems [12, 42]. Besides, sufficiently high
viscosity is essential for the biomaterial suspension to
overcome surface tension-driven droplet formation and be
drawn in form of straight filaments. High viscosity, on the
other hand, triggers clogging inside the nozzle tip and should
be optimized considering the diameter of the nozzle tip.
Considering prolonged fabrication time for printing scale-up
tissues and organs, one of the important disadvantages of
encapsulating living cells in biomaterials is that cellbiomaterial suspension needs to be stored considerable time in
the material reservoir that compromise cell viability and limits
their bioactivity. Thus, a more automated way of loading and
ejecting cell-biomaterial suspension is required for scale-up
tissue and organ fabrication. Recently, Novogen MMX
BioprinterTM [43] developed the technology that the printer
head has both suction and ejection capability enabling
automatic loading of suspension through the nozzle tip
occasionally instead of loading it manually prior to fabrication.
Mechanical strength and structural integrity of the fabricated
structures is a common drawback among bioprinting
techniques, which mainly use hydrogels due to high water
content and biocompatibility that allows permeation of
nutrients into and cellular products out of the gel [44]. Water
content increases their biocompatibility; however, it
deteriorates mechanical properties and processability
significantly. Although hydrogels possess unique properties,
they are intrinsically weak due to high water content and do
not withstand mechanical loading during and after gelation
process. Gelation of cell encapsulated hydrogels is a
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
crosslinking reaction initiated by a light, a chemical or thermal
transitions [45]. Photo-crosslinking processes compromise cell
viability and pose significant limitations on encapsulation of
cells within hydrogels [33]. Crosslinking through thermal
transitions limit applicability of hydrogels. Thermal transition
initiated crosslinking is not easy to handle after the process
while temperature changes can result in rapid degradation of
the printed thermogel that does not support cell viability in cell
media culture [46]. Chemical crosslinking can compromise
cell viability if any abrupt pH changes are observed; however,
non-acid involved crosslinking, i.e. sodium alginate, occurs
gently under mild conditions and at room temperature without
producing any toxic components, which has a great potential
for tissue engineering [47]. So, new hydrogels should be
tailored to enhance mechanical properties and processability
for specific bioprinting techniques towards advanced tissue
and organ fabrication. In addition, these materials should have
the ability to withstand sterilization while sterile conditions
certainly need to be acquired for process safety.
In general, cell encapsulation in biomaterial allows cell
patterning that has a great potential for direct organ printing;
however, subsequent extracellular matrix (ECM) formation,
digestion, and degradation of biomaterial matrix and
proliferation of encapsulated cells are not trivial to control.
There are still intrinsic limitations for bioprinting due to
limited cell proliferation and colonization while cells are
immobilized within hydrogels, and do not spread, stretch, and
migrate to generate the new tissue. Most recently, a new
concept was introduced by Mironov and his coworkers [48]
and has great potential in overcoming issues with cell
encapsulation within hydrogels. They proposed tissue
spheroids as building blocks for organ printing, where tissue
spheroids direct self-assembly toward organ fabrication.
Although tissue spheroids have been around for a long time in
demonstrating cellular fusion and tissue formation, where
hang-drop culture is one of the most common techniques [49],
bioprinting and assembling them in hydrogel media has
brought a significant potential for tissue engineering.
A
B
Fig. 3: (A) Hanging drop cultures with 20,000 chondrocytes
showing (B) spheroid fusion 7 days after harvesting, creating a
larger scale tissue.
Cell aggregates or cell-laden hydrogels as building blocks at
the microscale are selectively deposited to create larger tissues
[48, 50, 51]. In this approach, these “minitissues” are
considered as a biomaterial with certain measurable and
controllable properties and are assembled into tissues with
specific features through layer-by-layer stacking [52] or direct
4
assembly [48, 53]. Figure 3 shows hanging drop cultures,
where each spheroid contains 20,000 chondrocytes, and fusion
between spheroids was observed seven days after harvesting,
resulting in a larger-scale cartilage tissue. This shows a great
promise in obtaining larger-scale cartilage formation, where
researchers have already demonstrated regeneration of
cartilage tissue on rabbit knees with comparable properties
with that of natural cartilage tissue [42]. Tissue spheroids in
that study eliminated or minimized inclusion of biomaterials
and hence tissue regeneration achieved without in need of
scaffolds. In general, this also reduces complications related
with degradation of biomaterials and resulted toxic
byproducts. In addition, large scale tissues can be easily
obtained by fusion process, where cells in cell/laden hydrogels
cannot easily fuse [48]. By taking advantage of tissue
spheroids, the time needed for tissue maturation can be
significantly reduced, since each tissue spheroid contains a
large number of cells that can be printed at once. Moreover,
cell viability is enhanced due to the large cell seeding density
and less mechanical stress experienced compared with direct
cell manipulation [48]. Researchers have been investigating
means to reduce the actual size of tissue spheroids which is
around 500μm because the current size restricts
biomanufacturing of smaller-scale features considering the
average size of tissue spheroid as the resolution of the
fabrication technology. In addition, more efficient and
economical ways of fabricating tissue spheroids are also under
investigation. Iwasaki et al. introduced a new technique using
micro-patterned tissue culture plates for mass fabrication of
spheroids that has the potential for rapid scale-up organ
fabrication technology [54].
Fig. 4: Tissue spheroids for blood vessel printing: (a) deposition of
straight filaments containing a string of tissue spheroids (stained in
white) with agarose filaments as support material (stained in blue)
both around cellular filaments and inside the core, (b) design for
multicellular assembly with (c) printed samples with human
umbilical vein smooth muscle cells and human skin fibroblast cells
(reproduced from [55], image courtesy of Elsevier).
By using tissue spheroids, Forgacs and his coworkers at the
University of Missouri, Columbia, used RP-based bioprinting
together with multicellular spheroids made from smooth
muscle cells and fibroblasts, producing scaffold-free vascular
constructs [55]. Figure 4 illustrates vascular construct printing,
where tissue spheroids are printed sequentially in cylindrical
filaments from the bottom up. Upon fusion of tissue spheroids
followed by a tissue maturation process of three days postprinting, the support material is pulled away manually to
generate the lumen. Multiple cell types, including human
umbilical vein smooth muscle cells and human skin fibroblast
cells, were printed together to fabricate multicellular
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
constructs. The bioprinting platform used in this study, the
Novogen MMX BioprinterTM [43], has been recently
commercialized and specializes in developing bioprinting
across a broad array of cell types to create functional 3D
tissues. When building large-scale organ structures, the
mechanical integrity of the printed structures as well as the
integration of the vascular network with the rest of the organ
seems to be the major challenges to expanding the technology
for further applications.
III. ORGAN PRINTING AND ITS CHALLENGES
Organ printing is a computer-aided process in which cells
and/or cell-laden biomaterials are placed in the form of
aggregates, which then serve as building blocks and are further
assembled into a 3D functional organ. It’s an automated
approach that offers a pathway for scalable, reproducible mass
production of engineered living organs where multiple cell
types can be positioned precisely to mimic their natural
counterparts. Developing a functional organ requires advances
in and integration of three types of technology [56]: (1) cell
technology, which addresses the procurement of functional
cells at the level needed for clinical applications, (2)
biomanufacturing technology, which involves combining the
cells with biomaterials in a functional 3D configuration, and
(3) technologies for in vivo integration, which addresses the
issue of biomanufactured construct immune acceptance, in
vivo safety and efficacy, and monitoring of construct integrity
and function post-implantation. Success in fabrication of
functional organs highly depends on advancements in stem cell
technology. Stem cells, which are found in several tissues in
the human body [57], can self-renew to produce more stem
cells and differentiate into diverse specialized cell types to
form various organs [58]. A variety of cell types can be used
for this application, such as embryonic stem cells (ESCs) [57],
adult stem cells (ASCs) [59], most recently, induced
pluripotent stem (iPS) cells [60], and tissue-specific cell lines
[61, 62]. Although a patient’s stem cells can be differentiated
into organ-specific cells for organ printing, there is still risk of
tissue rejection by the receiver [56]. Stem cell behaviors can
even change during the bioprinting process. In addition, organ
fabrication necessitates various types of organ-specific cells,
which is not currently feasible considering the current isolation
and differentiation technologies. Although stem cells offer
great promise as an unlimited source of cells, a greater
understanding of and control over the differentiation process is
required in order to generate expandable organ-specific cells
in consistent quality with the desired phenotype. In this way,
rejection by the recipient side will be minimized posttransplantation. Moreover, imaging modalities such as
computed tomography (CT), positron emission tomography
(PET), and nuclear magnetic resonance (NMR) imaging
should be used to monitor the transplanted organ
noninvasively; NMR offers a unique advantage in monitoring
organ integrity and function without the need to modify the
cells genetically [63]. Although cell and transplantation
technology plays a crucial role in organ fabrication, this review
focuses on biomanufacturing technology, and the rest of the
5
paper discusses major challenges in the context of bioprinting
towards organ fabrication.
Despite the progress in tissue engineering, several
challenges must be addressed for organ printing to become a
reality. The most critical challenge in organ printing is the
integration of a vascular network, which is also a problem that
the majority of tissue engineering technologies are facing.
Without vascularization, engineered 3D thick tissue or organs
cannot get enough nutrients, gas exchange, and waste removal,
all of which are needed for maturation during perfusion. This
results in low cell viability and malfunction of artificial organs.
Systems must be developed to transport nutrients, growth
factors, and oxygen to cells while extracting metabolic waste
products such as lactic acid, carbon dioxide, and hydrogen
ions so the cells can grow and fuse together, forming the
organ. Cells in a large 3D organ structure cannot maintain their
metabolic functions without vascularization, which is
traditionally provided by blood vessels. Bioprinting
technology, on the other hand, currently does not allow multiscale tissue fabrication where bifurcated vessels are required to
be manufactured with capillaries to mimic natural vascular
anatomy. Although several researchers have investigated
developing vascular trees using computer models [64], only a
few attempts have been made toward fabricating bifurcated or
branched channels [55]. Successful maturation towards
functional mechanically integrated bifurcated vessels is still a
challenge.
In order to closely mimic natural organs, 3D vascularized
organs need to be fabricated using heterocellular aggregates,
which was discussed extensively in the literature [48]. In that
study, intraorgan vascular branched networks are envisioned to
be printed and maturated with the rest of the organ. Instead of
biomimetically designing and fabricating a vascular tree,
which seems to be one of the impediments down the road for
organ printing, we alternatively propose printable
semipermeable microfluidic channels to mimic a vascular
network in perfusing media and facilitating oxygenation for
cell viability, coaxing tissue maturation and formation.
Figure 5 illustrates our conceptual model of organ printing
through integration of vessel-like micro-fluidic channels with
cellular assembly, which has been featured in the literature [1,
65]. The concept allows constructing structures with microfluidic channels acting like blood vessels, allowing media
perfusion through an extracellular matrix that can support cell
viability in 3D. In Fig. 5, vessel-like micro-fluidic channels are
printed in a 0°-90° lay-down pattern to develop 3D structures.
Oxygenized perfusion media can be pumped into channels for
circulation purposes. Round ends are considered for the zigzag
deposition pattern to prevent blockage of media flow. After
parallel printing of each micro-fluidic channel layer, cellular
spheroids, i.e., tissue spheroids or cell encapsulated
microspheres, can be deposited between fluidic channels in the
form of droplets using another robotic arm; semipermeable
fluidic channels allow transport and diffusion of media to the
cellular environment. The proposed strategy enables printing
semi-permeable micro-fluidic channels in tandem with printing
cellular assembly. Concurrent printing has the potential to
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
reduce the fabrication time, which is crucial for the
development of scale-up technologies (see Fig. 7(C)).
Fig. 5: Concept of 3D organ-printing technology where vessel-like
microfluidic channels are printed in tandem with tissue spheroids
layer by layer. The printed structure can be then connected to a
bioreactor for media perfusion (figures are not to scale).
Advantages of tissue spheroids have already been discussed
in Section 2; spheroids significantly enhance cell viability of
stem-cell-derived organ-specific cells, allowing the
development of further organ fabrication techniques in the
near future. Printing encapsulated cells in spheroids also
greatly reduces shear-stress-induced cell damage compared to
printing cells directly loaded within biomaterial media. In
general, cells are subjected to shear stress during the printing
process, and cell injury and DNA damage should be
minimized. The proposed concept can also allow inclusion of
multiple cell types in a spatially organized way by integrating
another printer unit mounted on the robotic arm to print a
secondary type of spheroids precisely.
6
printed microfluidic channels showed 97.6±1.2% one day
post-printing and maintained high cell viability on day 4, with
a percentage of 95.8±1.2 where living cells are shown around
the lumen on the cut-away view (see Fig. 6 (C)).
It should be ensured that the printed organ functions
correctly. For example, a bioprinted pancreatic organ must be
able to produce and secrete insulin just like its natural
counterpart. Thus, multiple organ-specific cell types are
required to be spatially organized to form the complex
architecture of an organ. Advancements in bioprinting
processes to precisely and selectively place cells, are
promising when it comes to achieving heterocellular
architectures. Figure 7 illustrates some of the bioprinters that
can facilitate printing heterocellular structures, including
EnvisionTEC 3D Bioplotter® (where solid filaments are
printed in a secondary plotting media), Novogen MMX
BioprinterTM (which allows printing heterocellular tissue
aggregates and hydrogels as support material), and the MultiArm BioPrinter (MABP) (which can print tissue spheroids in
tandem with vessel-like microfluidic channels). These
technologies can be applied for cellular aggregate
biofabrication in such a way as to facilitate self-assembly of
the tissue spheroids to mimic the natural process of tissue and
organ morphogenesis.
A
B
C
Fig. 6: Printed cellular microfluidic channels: (A) perfusion of
oxygenized cell type media, (B) media flow with intentionally
generated air bubbles and (C) the laser confocal image of the midplane showing single lumen channel with live/dead staining where
CPCs are labeled with calcein AM and ethidium homodimer.
Figure 6 illustrates oxygenized media perfusion through a
printed cellular micro-fluidic channel 44cm in length and
1mm wide with a 500µm lumen diameter. The cell media is
circulated through the channel without any blockage or
swirling, which shows a great potential for developing
embedded channels and serving as a vascular network for
thick tissue fabrication. Direct printing of these channels
allows them to be integrated within a hybrid bioprinting
platform and also facilitates patterning them into very
complex shapes. Currently, mechanobiological properties of
printed channels are under investigation, and electrospun
nanofiber reinforcement is performed to match the mechanical
properties with that of blood vessels. Elasticity and tensile
strength are essential to be biomimetically mediated. Viability
of cartilage progenitor cells (CPCs) encapsulated within
Fig. 7: Bioprinters: (A) 3D Bioplotter® (designed by envisionTEC
GmbH, Gladbeck, Germany [66]), (B) Novogen MMX BioprinterTM
(designed by Organovo, San Diego, California, USA [43]) and (C)
Multi-Arm Bioprinter-“MABP” (designed by the University of Iowa,
Iowa City, Iowa, USA).
In order to fabricate scalable organs such as a mouse liver
which has 1.3x108 cells per gram [67], the bioprinter needs to
run for several hours considering the resolution of the system
as in microscale. During prolonged fabrication time, it should
deliver biological substances through micro-nozzle or other
means without clogging problems or collision issues between
the nozzle tip and the printed structure. In addition,
sterilization is also crucial during bioprinting and necessitates
construction of bioprinters on a small scale to fit in standard
biosafety cabinets. In general, commercial bioprinters can cost
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
substantially high around $100-200k depending on their
unique capabilities, where home-made bioprinters on the other
hand roughly cost less than quarter of their commercial
counterparts.
Another challenging site for organ printing technology is the
rapid or accelerated tissue maturation process, where printed
organ constructs should be rapidly fused, remodeled, and
maturated toward a solid construct, ensuring mechanical
rigidity for transplantation. Collagen and elastin are some of
the proteins in the extracellular matrix of human organs, and
they are abundant in the connective tissue stromal of
parenchymal organs [68, 69]. Thus, production and deposition
of these proteins during tissue maturation are essential to
enhancing the mechanical properties of the printed organs.
After the organ-printing process, the fabricated structure needs
to be transferred to the bioreactor. Bioreactors such as an
irrigation dripping perfusion bioreactor can be used to
expedite the tissue maturation and organ formation process,
providing an optimal environment. The transfer of a printed
organ to the bioreactor should be performed gently without
inducing any vibration that can degenerate the integrity of
fragile gel-based structures. Thus, future bioprinters can be
enclosed in a bioreactor system that will allow direct and rapid
connection of printed structures to perfusion channels.
IV. PROPOSED ENVISIONED FUTURE
Considering the pathway from isolation of stem cells to
transplantation into a human, seamlessly automated protocols
and systems are essential for customized functional organ
fabrication. This pathway includes (i) blueprint modeling of an
organ with its vascular architecture, (ii) generation of a process
plan for bioprinting, (iii) isolation of stem cells, (iv)
differentiation of stem cells into organ-specific cells, (v)
preparation and loading of organ-specific cells and blood
vessel cells as well as support medium, and (vi) bioprinting
process followed by organogenesis in a bioreactor for
transplantation. In order to accelerate the entire process, one
can envision a bioreactor enclosing a bioprinter that will
immediately facilitate fusion of tissues and keep the printed
organ until desired maturation is achieved. In this way, a
customized and automated system will facilitate delivering
artificial organs to transplant patients in a reasonable amount
of time, preferably at earlier stages of their diseases when
patients are healthier so that they are better able to withstand
surgical interventions.
Miniature organs can be considered a future trend in organ
printing and might be a transition towards fully functioning
organs. Miniature organs, which can also be called a “factory
in the human body”, can be built in smaller scale than their
natural counterparts and closely perform the most vital
function of the associated organ, such as a pancreatic organ
that can produce and secrete insulin in substantial amounts to
regulate the glucose level to normoglycemia in the human
body. The miniature organ can be transplanted and placed in
less immune-responsive sites in the human body as an
extravascular device or attached to blood vessels as an
7
intravascular device to secrete insulin to the bloodstream.
Although the pancreas serves two functions, which are carried
out by two different cell groups within the organ, a patient
with diabetes will be interested in correcting hyperglycemia
through fabrication and transplantation of a pancreatic organ
that can restore the function of the endocrine portions of the
pancreas, which makes only about 2% of pancreas cells. The
endocrine portion is made of approximately a million cell
clusters called islets of Langerhans [70]. Four main cell types
exist in the islets. α, β, delta and gamma cells secrete glucagon,
insulin, somatostatin (regulates/stops α and β cells) and
pancreatic polypeptide, respectively. Miniature organs can also
be designed and fabricated to bring new functionalities and
superiorities in the human body rather than restoring the
functionality of their natural counterparts, such as living
organs that can continuously generate electricity to eliminate
the use of batteries for internal devices such as peace makers.
Another future trend in organ printing is in-situ printing,
where living organs can be printed in the human body during
operations. Currently, in-situ bioprinting has already been
tested for repairing external organs such as skin [38], where
the wounded section is filled with multiple cells, including
human keratinocytes and fibroblast, with stratified zones
throughout the wound bed. A transitional approach seems to
be logical to advance the state of the art through printing and
repairing partially damaged, diseased, or malfunctioning
internal organs that do not have self-repair characteristics, such
as the liver. With the recent advancements in robot-assisted
surgery, computer-controlled robotic bioprinters will lead the
evolution of this technology in the very near future.
V. CONCLUSIONS
This paper discusses the current state of the art in
bioprinting with recent trends in organ printing technology
towards fabrication of living organs for transplant, where
prototype organs can be developed using layer-by-layer
technology in the very near future. Although the technology
shows a great deal of promise, there is still a long way to go to
practically realize this ambitious vision. Overcoming current
impediments in cell technology, biomanufacturing technology,
and technologies for in-vivo integration is essential for
developing seamlessly automated technology from stem cell
isolation to transplantation.
ACKNOWLEDGMENT
The authors would like to thank A. Lehman from the
Ignacio Ponseti Orthopaedic Cell Biology Lab (The University
of Iowa) for the fabrication of chondrocyte spheroids.
REFERENCES
[1]
[2]
[3]
H. Chen and I. Ozbolat. (2013) Manufacturing Living Things.
Industrial Engineer Magazine, vol. 45, 30-34.
R. Langer and J. P. Vacanti, "Tissue engineering," Science, vol.
260, pp. 920-6, 1993.
M. Sittinger, J. Bujia, W. W. Minuth, C. Hammer, and G. R.
Burmester, "Engineering of cartilage tissue using bioresorbable
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
[4]
[5]
[6]
[7]
[8]
[9]
[10]
[11]
[12]
[13]
[14]
[15]
[16]
[17]
[18]
[19]
[20]
[21]
[22]
[23]
[24]
[25]
polymer carriers in perfusion culture," Biomaterials, vol. 15, pp.
451-6, 1994.
J. R. Porter, T. T. Ruckh, and K. C. Popat, "Bone tissue
engineering: a review in bone biomimetics and drug delivery
strategies," Biotechnol Prog, vol. 25, pp. 1539-60, 2009.
S. Korossis, F. Bolland, E. Ingham, J. Fisher, J. Kearney, and J.
Southgate, "Review: tissue engineering of the urinary bladder:
considering structure-function relationships and the role of
mechanotransduction," Tissue Eng, vol. 12, pp. 635-44, 2006.
S. V. Nolte, W. Xu, H. O. Rennekampff, and H. P. Rodemann,
"Diversity of fibroblasts--a review on implications for skin tissue
engineering," Cells Tissues Organs, vol. 187, pp. 165-76, 2008.
A. Atala, "Tissue engineering for bladder substitution," World J
Urol, vol. 18, pp. 364-70, 2000.
C. Barnatt, "Organ Printing Concept," Bioprinter_Holdout, Ed.,
ed: www.ExplainingTheFuture.com, 2011.
J. A. Barron, P. Wu, H. D. Ladouceur, and B. R. Ringeisen,
"Biological laser printing: a novel technique for creating
heterogeneous 3-dimensional cell patterns," Biomed Microdevices,
vol. 6, pp. 139-47, 2004.
T. Boland, T. Xu, B. Damon, and X. Cui, "Application of inkjet
printing to tissue engineering," Biotechnol J, vol. 1, pp. 910-7,
2006.
V. Mironov, "Printing technology to produce living tissue," Expert
Opin Biol Ther, vol. 3, pp. 701-4, 2003.
F. P. W. Melchels, M. A. N. Domingos, T. J. Klein, J. Malda, P. J.
Bartolo, and D. W. Hutmacher, "Additive manufacturing of tissues
and organs," Progress in Polymer Science, vol. 37, pp. 10791104, 2012.
H. Lipson and M. Kurman, Fabricated: The New World of 3D
Printing Wiley, 2013.
D. J. Odde and M. J. Renn, "Laser-guided direct writing for
applications in biotechnology," Trends Biotechnol, vol. 17, pp.
385-9, 1999.
A. Ovsianikov, M. Gruene, M. Pflaum, L. Koch, F. Maiorana, M.
Wilhelmi, et al., "Laser Printing of Cells into 3D Scaffolds,"
Biofabrication, vol. 2, p. 014104, 2010.
Y. Nahmias, R. E. Schwartz, C. M. Verfaillie, and D. J. Odde,
"Laser-guided direct writing for three-dimensional tissue
engineering," Biotechnol Bioeng, vol. 92, pp. 129-36, 2005.
X. Cui, K. Breitenkamp, M. G. Finn, M. Lotz, and D. D. D'Lima,
"Direct human cartilage repair using three-dimensional bioprinting
technology," Tissue Eng Part A, vol. 18, pp. 1304-12, 2012.
H. C.L. and Y. Huang, "Alginate microsphere fabrication using
bipolar wave-based drop-on-demand jetting," Journal of
Manufacturing Processes, vol. 14, pp. 98-106, 2012.
C. Xu, W. Chai, Y. Huang, and R. Marwald, "Scaffold-free inkjet
printing of three-dimensional zigzag cellular tubes,"
Biotechnology and Bioengineering, vol. 209, pp. 3152-3160,
2012.
C. L. Herran and Y. Huang, "Alginate microsphere fabrication
using bipolar wave-based drop-on-demand jetting," Journal of
Manufacturing Processes, vol. 14, pp. 98-106, 2012.
C. Xu, W. Chai, Y. Huang, and R. R. Markwald, "Scaffold-free
inkjet printing of three-dimensional zigzag cellular tubes,"
Biotechnol Bioeng, vol. 109, pp. 3152-60, 2012.
X. Cui and T. Boland, "Human microvasculature fabrication using
thermal inkjet printing technology," Biomaterials, vol. 30, pp.
6221-7, 2009.
L. Weiss, C. Amon, S. Finger, E. Miller, and D. Romero,
"Bayesian computer-aided experimental design of heterogenous
scaffolds for tissue engineering," Computer-Aided Design, vol. 37,
pp. 1127-1139, 2005.
P. G. Campbell and L. E. Weiss, "Tissue Engineering with the aid
of inkjet printers," Expert Oprinion on Biological Therapy, vol. 7,
pp. 1123-1127, 2007.
K. Jakab, C. Norotte, F. Marga, K. Murphy, G. VunjakNovakovic, and G. Forgacs, "Tissue engineering by self-assembly
and bio-printing of living cells," Biofabrication, vol. 2, p. 022001,
2010.
[26]
[27]
[28]
[29]
[30]
[31]
[32]
[33]
[34]
[35]
[36]
[37]
[38]
[39]
[40]
[41]
[42]
[43]
[44]
[45]
[46]
[47]
8
Y. Yan, X. Wang, Y. Pan, H. Liu, J. Cheng, Z. Xiong, et al.,
"Fabrication of viable tissue-engineered constructs with 3D cellassembly technique," Biomaterials, vol. 26, pp. 5864-71, 2005.
R. Chang, J. Nam, and W. Sun, "Direct Cell Writing of 3D Microorgan for In Vitro Pharmacokinetic Model," Tissue Engineering,
vol. 14, pp. 157-169, 2008.
S. Khalil and W. Sun, "Biopolymer deposition for freeform
fabrication of hydrogel tissue constructs," Material Science and
Engineering: C, vol. 27, pp. 469-478, 2007.
S. Khalil, F. Nam, and W. Sun, "Multi-nozzle deposition for
construction of 3-D biopolymer tissue scaffolds," Rapid
Prototyping Journal, vol. 11, pp. 9-17, 2005.
K. Nair, K. Yan, and W. Sun, "A Multi-Level Numerical Model
for Qunatifying Cell Deformation in Encapsulated Alginate
Structures " Journal of Mechanics of Materials and Structures,
vol. 6, pp. 1121-1139, 2007.
N. N. Kachouie, Y. Du, H. Bae, M. Khabiry, A. F. Ahari, B.
Zamanian, et al., "Directed assembly of cell-laden hydrogels for
engineering functional tissues," Organogenesis, vol. 6, pp. 23444, 2010.
N. R. Schiele, D. B. Chrisey, and D. Corr, "Gelatin-Based Laser
Direct-Write Technique for the Precise Spatial Patterning of Cells
" Tissue Eng Part C, vol. 17, pp. 289-298, 2011.
T. Billiet, M. Vandenhaute, J. Schelfhout, S. Van Vlierberghe, and
P. Dubruel, "A review of trends and limitations in hydrogel-rapid
prototyping for tissue engineering," Biomaterials, vol. 33, pp.
6020-41, 2012.
S. Catros, B. Guillotin, M. Bačáková, J.-C. Fricain, and F.
Guillemot, "Effect of laser energy, substrate film thickness and
bioink viscosity on viability of endothelial cells printed by LaserAssisted Bioprinting," Applied Surface Science, vol. 257, pp.
5142-5147, 4/1/ 2011.
J.-W. Choi, H.-C. Kim, and R. Wicker, "Multi-material
stereolithography," Journal of Materials Processing Technology,
vol. 211, pp. 318-328, 2011.
B. Guillotin and F. Guillemot, "Cell patterning technologies for
organotypic tissue fabrication," Trends in Biotechnology, vol. 29,
pp. 183-190, 2011.
T. Xu, C. Baicu, M. Aho, M. Zile, and T. Boland, "Fabrication
and characterization of bio-engineered cardiac pseudo tissues,"
Biofabrication, vol. 1, p. 035001, 2009.
K. W. Binder, "In Situ Bioprinting of the Skin," PhD, Molecular
Genetics and Genomics, Wake Forest University, Winston-Salem,
North Carolina, 2011.
X. Cui, D. Dean, Z. M. Ruggeri, and T. Boland, "Cell damage
evaluation of thermal inkjet printed Chinese hamster ovary cells,"
Biotechnol Bioeng, vol. 106, pp. 963-9, 2010.
T. Xu, J. Jin, C. Gregory, J. J. Hickman, and T. Boland, "Inkjet
printing of viable mammalian cells," Biomaterials, vol. 26, pp. 9399, 2005.
K. Nair, M. Gandhi, S. Khalil, K. C. Yan, M. Marcolongo, K.
Barbee, et al., "Characterization of cell viability during bioprinting
processes," Biotechnology Journal, vol. 4, pp. 1168-1177, 2009.
J. Lee, M. Sato, H. Kim, and J. Mochida, "Transplantation of
scaffold-free spheroids composed of synovium-derived cells and
chondrocytes for the treatment of cartilage defects of the knee,"
Europ. Cells and Materials, vol. 22, pp. 275-290, 2011.
OrganovoTM. (December 7). Available: http://www.organovo.com/
G. Y. Huang, L. H. Zhou, Q. C. Zhang, Y. M. Chen, W. Sun, F.
Xu, et al., "Microfluidic hydrogels for tissue engineering,"
Biofabrication, vol. 3, p. 012001, 2011.
A. S. Hoffman, "Hydrogels for biomedical applications,"
Advanced Drug Delivery Reviews, vol. 64, pp. 18-23, 2012.
N. Fedorovich, J. Wijn, A. Verbout, J. Alblas, and W. Dhert,
"Three-dimensional fiber deposition of cell-laden, viable,
patterned constructs for bone tissue printing," Tissue Eng Part A,
vol. 14, pp. 127-133, 2008.
J. L. Drury and D. J. Mooney, "Hydrogels for tissue engineering:
scaffold design variables and applications," Biomaterials, vol. 24,
pp. 4337-4351, 2003.
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].
This article has been accepted for publication in a future issue of this journal, but has not been fully edited. Content may change prior to final publication.
TBME-01840-2012
[48]
[49]
[50]
[51]
[52]
[53]
[54]
[55]
[56]
[57]
[58]
[59]
[60]
[61]
[62]
[63]
[64]
[65]
[66]
[67]
[68]
[69]
[70]
V. Mironov, Visconti, R.P. ,Kasyanov V., Forgacs G., Drake C.J.,
Markwald, R. R., "Organ printing: Tissue spheroids as building
blocks," Biomaterials, vol. 30, pp. 2164-2174, 2009.
L. Chopinet, L. Wasungu, and M.-P. Rols, "First explanations for
differences in electrotransfection efficiency in vitro and in vivo
using spheroid model," International Journal of Pharmaceutics,
vol. 423, pp. 7-15, 2012.
X. Yang, V. Mironov, and Q. Wang, "Modeling fusion of cellular
aggregates in biofabrication using phase field theories " Journal of
Theoretical Biology, vol. 303, pp. 110-118, 2012.
V. Mironov, V. Kasyanov, and R. Markwald, "Organ printing:
from bioprinter to organ biofabrication line," Current Opinion in
Biotechnology, vol. 22, pp. 667-673, 2011.
N. L'Heureux, T. N. McAllister, and L. M. de la Fuente, "Tissueengineered blood vessel for adult arterial revascularization," N
Engl J Med, vol. 357, pp. 1451-3, 2007.
Y. Du, E. Lo, S. Ali, and A. Khademhosseini, "Directed assembly
of cell-laden microgels for fabrication of 3D tissue constructs,"
Proc Natl Acad Sci U S A, vol. 105, pp. 9522-7, 2008.
A. Iwasaki, T. Matsumoto, G. Tazaki, H. Tsuruta, H. Egusa, H.
Miyajima, et al., "Mass Fabrication of Small Cell Spheroids by
Using Micro-patterned Tissue Culture Plate," Advanced
Engineering Materials, vol. 11, pp. 801-804, 2009.
C. Norotte, F. S. Marga, L. E. Niklason, and G. Forgacs,
"Scaffold-free vascular tissue engineering using bioprinting,"
Biomaterials, vol. 30, pp. 5910-7, 2009.
R. Lanza, R. Langer, and J. Vacanti, Principles of Tissue
Engineering, 3 ed.: Elsevier, 2007.
W. W. Thein-Han and Y. Kitiyanant, "Chitosan scaffolds for in
vitro buffalo embryonic stem-like cell culture: an approach to
tissue engineering," J Biomed Mater Res B Appl Biomater, vol.
80, pp. 92-101, 2007.
R. L. Gardner, "Stem cells: potency, plasticity and public
perception," J Anat, vol. 200, pp. 277-82, 2002.
R. S. Tuan, G. Boland, and R. Tuli, "Adult mesenchymal stem
cells and cell-based tissue engineering," Arthritis Res Ther, vol. 5,
pp. 32-45, 2003.
M. T. Lam and M. T. Longaker, "Comparison of several
attachment methods for human iPS, embryonic and adiposederived stem cells for tissue engineering," J Tissue Eng Regen
Med, 2012.
D. Seol, D. J. McCabe, H. Choe, H. Zheng, Y. Yu, K. Jang, et al.,
"Chondrogenic progenitor cells respond to cartilage injury,"
Arthritis Rheum, vol. 64, pp. 3626-37, 2012.
Y. Yu, "Identification and characterization of cartilage progenitor
cells by single cell sorting and cloning," University of Iowa, 2012.
C. L. Stabler, R. C. Long, I. Constantinidis, and A. Sambanis, "In
vivo noninvasive monitoring of a tissue-engineered construct
using 1H-NMR spectroscopy. ," Cell Transplant, vol. 14, pp. 139149, 2005.
W. L. Mondy, D. Cameron, J.-P. Timmermans, N. D. Clerk, A.
Sasov, C. Casteleyn, et al., "Computer-aided design of
microvasculature systems for use in vascular scaffold production,"
Biofabrication, vol. 1, p. 035002, 2009.
J. Thilmany. (2012) Printed Life. Mechanical Engineering
Magazine, vol. 134, pp. 44-47.
envisionTEC. (December 7). Available: www.envisiontec.com
R. Marcos, R. A. Monteiro, and E. Rocha, "Design-based
stereological estimation of hepatocyte number, by combining the
smooth optical fractionator and immunocytochemistry with anticarcinoembryonic antigen polyclonal antibodies," Liver Int, vol.
26, pp. 116-24, 2006.
B. D. McLees, G. Schleiter, and S. R. Pinnell, "Isolation of type III
collagen from human adult parenchymal lung tissue,"
Biochemistry, vol. 16, pp. 185-90, 1977.
V. Richmond, "Lung parenchymal elastin isolated by nondegradative means," Biochim Biophys Acta, vol. 351, pp. 173-7,
1974.
G. E. Hellman B, Grapengiesser E, Dansk H, Salehi A, "Insulin
oscillations--clinically important rhythm. Anti-diabetics should
increase the pulsative component of the insulin release,"
Lakartidningen, pp. 32-33, 2007.
9
Ibrahim T. Ozbolat received dual B.S.
degrees in Industrial Engineering and
Mechanical Engineering from the Middle
East Technical University, Ankara,
Turkey, in 2006 and 2007, respectively,
and the Ph.D. degree in industrial
engineering from University at Buffalo,
New York, USA, in 2011.
He is an Assistant Professor of the Department of
Mechanical and Industrial Engineering and a Research Faculty
at the Center for Computer-Aided Design at The University of
Iowa, Iowa City, IA, USA. His research interests are
biomanufacturing, tissue engineering, manufacturing and
design, virtual manufacturing, computer-aided design and
electronics manufacturing.
Prof. Ozbolat belongs to IIE, ASME, APM and TERMIS.
Yin Yu received his M.D. degree from
Medical College of Nantong University,
China, in 2010 and his M.S. degree from
Biomedical Engineering at the University,
Iowa City, Iowa, USA.
His research mainly focuses on stem
cell biology, tissue engineering and
biomanufacturing.
He is a member of ASME, TERMIS, BMES and ICMRS.
Copyright (c) 2013 IEEE. Personal use is permitted. For any other purposes, permission must be obtained from the IEEE by emailing [email protected].