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Cell and organ printing 2Fusion of cell aggregates in three-dimensional gels.

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Cell and Organ Printing 2:
Fusion of Cell Aggregates in
Three-Dimensional Gels
Department of Bioengineering, Clemson University, Clemson, South Carolina
Department of Cell Biology and Anatomy, Medical University of South Carolina,
Charleston, South Carolina
Pacific Northwest National Laboratory, Richland, Washington
We recently developed a cell printer (Wilson and Boland, 2003) that
enables us to place cells in positions that mimic their respective positions in
organs. However, this technology was limited to the printing of two-dimensional (2D) tissue constructs. Here we describe the use of thermosensitive
gels to generate sequential layers for cell printing. The ability to drop cells
on previously printed successive layers provides a real opportunity for the
realization of three-dimensional (3D) organ printing. Organ printing will
allow us to print complex 3D organs with computer-controlled, exact placing
of different cell types, by a process that can be completed in several minutes.
To demonstrate the feasibility of this novel technology, we showed that cell
aggregates can be placed in the sequential layers of 3D gels close enough for
fusion to occur. We estimated the optimum minimal thickness of the gel
that can be reproducibly generated by dropping the liquid at room temperature onto a heated substrate. Then we generated cell aggregates with the
corresponding (to the minimal thickness of the gel) size to ensure a direct
contact between printed cell aggregates during sequential printing cycles.
Finally, we demonstrated that these closely-placed cell aggregates could
fuse in two types of thermosensitive 3D gels. Taken together, these data
strongly support the feasibility of the proposed novel organ-printing technology. Anat Rec Part A 272A:497–502, 2003. © 2003 Wiley-Liss, Inc.
Key words: organ printing; cell printing; cell aggregates; cell
fusion; hydrogel
Several approaches are currently being used to address
the shortage of donor organs, including artificial mechanical organs, xenotransplantation (using animal organs),
tissue engineering, and regenerative medicine. Some artificial organs are already on the market, but they can
significantly diminish a patient’s quality of life and may
produce unwanted side effects. Xenotransplanation is a
promising approach, especially when it involves the use of
organs from transgenic animals with a reduced capacity to
induce acute immune response after transplantation, and
when it is combined with emerging methods of immunotolerance management. However, there is still great concern about the potential spreading of animal viruses (Cooper et al., 2002). Regenerative medicine, or the repair of
injured or diseased organs in vivo by gene therapy or cell
transplantation, is also a very promising and sophisti©
cated approach. However, gene therapy and cell transplantation, as well as drug therapy, are most effective in
the early stages of a disease. At terminal stages, or for
injured tissue, there is still a need for organ replacement.
Thus it is has been proposed that tissue engineering,
Grant sponsor: NASA EPSCoR Grant; Grant number: HEDS6.
*Correspondence to: Thomas Boland, Department of Bioengineering, Clemson University, Clemson, SC 29634. Fax: (864)
Received 28 March 2002; Accepted 17 February 2003
DOI 10.1002/ar.a.10059
which is the growing of organs in vitro from a patient’s
own cells, is a reasonable approach to address the shortage of donor organs.
The classical tissue engineering approach involves the
use of solid, rigid scaffolds from polyglycolic acid (PGA)
and isolated cells (Langer and Vacanti, 1993). It is based
on the premise that seeding cells in a bioreactor on porous
biodegradable scaffolds will be sufficient to generate organs. However, there are at least four problems with this
1. Cell penetration and seeding is not very effective. It
proceeds on the time scale of months and is not uniform
throughout the scaffold. Although significant progress
has been made in designing scaffolds that enable effective seeding and cell migration (Ma and Zhang, 2001),
it is still far from optimal.
2. Organs usually consist of many cell types, and the need
to place different cell types in specific positions is a very
challenging technical problem in solid scaffold design.
3. The rigid, solid scaffolds made from PLA are not optimal for engineering contractile tissue, such as heart
and vascular tubes.
4. The main problem with using solid scaffold seeding
technology (for constructs larger than 200 ␮) is the
absence of vascularization.
The lack of vascularization has been addressed by a
“rolling” approach, which has proven to be very effective
for building blood vessels (L’Heureux et al., 1998), but it
cannot be adapted for complex three-dimensional (3D)
organs. Embedding technology is also very promising (Nerem and Seliktar, 2001), but it can only be applied to
multilayer tubular organs, and also suffers from the absence of a direct placing mechanism.
Recently, rapid prototyping (RP) technology has been
used to fabricate physical models of hard tissues, tissue
scaffolds, and custom-made tissue implant prostheses
(Potamianos et al., 1998; Holck et al., 1999; Winder et al.,
1999). RP technology has been used to produce novel scaffolds with controllable porosity and channel sizes, potentially allowing for vacularization (Zein et al., 2002). Although the technique is versatile and can reproduce
anatomical structures (Sodian et al., 2002), its main limitation is the lack of polymers with suitable mechanical
properties for soft-tissue constructs. (Hutmacher et al.,
2001). This technique also has the drawback that cells
must be placed in exact positions in the 3D printed scaffold.
We recently developed a cell printer (Wilson and Boland,
2003) that enables us to place cells in positions that mimic
their positions in an organ. The printer can put up to nine
solutions of cells or polymers into a specific place by the
use of specially designed software, and print two-dimensional (2D) tissue constructs. However, up to now this
technology was limited to the printing of only 2D tissue
constructs. A new opportunity for extending the printing
technology to three dimensions is created by the use of
thermo-reversible gels (Gutowska et al., 2001). Nontoxic,
biodegradable, thermo-reversible gels, which are fluid at
20°C and gel above 32°C, are used as a sort of “paper” on
which tissue structures can be printed, and the cells are
the “ink.” Successive layers could be generated just by
dropping another layer of gel onto an already printed
surface (see Fig. 1).
Fig. 1. Principle of organ printing. [Color figure can be viewed in the
online issue, which is available at]
This technology, which we term “organ printing,” enables the printing of complex 3D organs, with exact placing of different cell types in 3D engineered organs, in only
a few minutes. To demonstrate the feasibility of this novel
technology, we showed that we can place cell aggregates in
a 3D gel closely enough for fusion to occur. To do this, we
estimated the minimal thickness of gel that can be reproducibly generated by dropping the liquid at room temperature onto heated substrata. Then, on the basis of this
experimentally estimated parameter, we designed cell aggregates with appropriate sizes to ensure a direct contact
between printed cell aggregates during sequential printing cycles. Finally, we showed that closely-placed cell aggregates could fuse in two types of 3D gels.
Cell Printer Preparation
A description of the printer is given elsewhere (Wilson
and Boland, 2003). Sterile stainless steel needles with luer
plastic hubs were kept on ice until they were filled with
the polymer gel of interest. They were then screwed into
the ethanol-sterilized luer fittings of the print head. Care
was taken to maintain a needle temperature below 4°C. It
is important to maintain a low temperature when working
with collagen because it will gel irreversibly and block the
needles at room temperature. The print head and the
printer were placed onto heated petri dishes. The dishes
were placed into specially designed aluminum blocks that
were kept at 36°C by a heater (VWR Scientific, Atlanta,
GA). The heater, blocks, dishes, and printer were put into
a sterile laminar flow hood.
Preparation of Thermosensitive Gels
A poly[N-isopropylacrylamide-co-2-(N,N-dimethylamino)ethyl acrylate] copolymer, denoted henceforth as K-70,
was used for the gel experiments. A molecular weight of
510 kD was estimated by gel permeation chromatography
(GPC) with a light-scattering detector. The polymer synthesis and characterization are described elsewhere (Gutowska et al., 2001).
A 10 wt % polymer solution in cold deionized water was
put into an external ice bath and stirred magnetically
overnight. After the polymer was completely dissolved,
the pH of the solution was adjusted to 7.0 with 0.1 M
NaOH, using accurate pH strips. The polymer solution
was sterilized for 30 min in an autoclave, cooled, and then
redissolved. The solutions were allowed to become completely clear and were mixed well before further use. The
polymer solutions were combined with an equal amount of
2⫻ cell culture medium consisting of Eagle’s minimum
medium (MEM) supplemented with 10% fetal bovine serum (FBS) and 1% antibiotic solution. The final polymer
concentration was not lower than 4 –5 wt %. To visualize
the polymer layers, half of the solutions were dyed by
adding 30 ␮l of 1% trypan blue dye to each milliliter of
polymer solution.
Preparation of Collagen Gels
Collagen gels were prepared according to the method of
Storm and Michalopoulous (1982), with the exception that
lyophilized calf skin collagen rather than rat tail collagen
was used. A 3.33 mg/ml collagen solution was prepared by
slowly dissolving the collagen in acetic acid (1%) at 0°C.
The gel was prepared by adding 156.5 ␮l of sterile, filtered
10⫻ DMEM and 30 ␮l of sterile 0.34N NaOH to each
milliliter of collagen solution. The resulting solutions were
kept on ice until they were used.
Estimation of Gel Thickness
The thickness of the gel layers was evaluated by depositing a series of drops onto the heated petri dishes. Each
drop was allowed to gel before a new, smaller drop was
placed on top of the gel. The series of staggered gel layers
thus created was placed under a microscope and visualized. The movement (in microns) of the microscope objective was measured as the focal plane shifted from one
layer to the next.
Preparation of Cell Aggregates
Bovine aortal endothelial cells (American Type Culture
Collection, Manassas, VA) (passage 10) were expanded in
T25 culture flasks in the presence of MEM supplemented
with 10% FBS and 1% antibiotic solution in a 5% CO2
incubator maintained at 36°C. The media were changed at
day 1 and subsequently every other day. After growing to
a confluent monolayer, the cells were washed by replacing
the media with isotonic phosphate-buffered saline (PBS).
They were then incubated in a PBS solution containing a
millimolar solution of the tripeptide arginine-glycine-aspartic acid (RGD) and agitated for 30 – 45 min in the
incubator. This caused the cells attached to the edges of
the flask to dislodge. Further mechanical scraping of the
flask with a sterile glass pipette caused the cells to detach
in aggregates. The aggregates were collected, centrifuged
at 1,000 rpm for 5 min, and resuspended in 0.5 ml of
media. The aggregates were then reseeded on collagen
and K-70 gels.
Live/Dead Assay
The viability of the cells was assessed with a commercially available live/dead assay (Molecular Probes Inc.,
Eugene, OR). The samples were rinsed with PBS, and
incubated for 30 min in a solution of calcein AM and
ethidium homodimer-1 in PBS according to the manufacturer’s protocol. Fluorescence was observed in an inverted
epifluorescent microscope (Nikon Diaphot 300, Nikon Inc.,
Melville, NY) using a DAPI/FITC/TRITC triple-band filter.
Alternate layers of clear and tryan blue K-70 gel were
analyzed under a microscope. Figure 2 shows a top view of
the layered gel. The thickness of the layers varied between
200 and 500 ␮. During the addition of the polymer solutions to the already gelled layers, it was obvious that only
minimal (if any) mixing of the layers occurred. In the
gelled state the polymer is hydrophobic, and one could
clearly observe that the liquid solution drops did not
spread out on the gel.
The size of the aggregates was also measured. We obtained aggregates with an average diameter of 540 ⫾ 183
␮m (n ⫽ 18). Figure 3 shows a representative image of the
endothelial aggregates used in this study.
Directed Fusion of Closely-Placed Cell
To prove that a directed fusion of aggregates could be
achieved, we performed an additional series of experiments using collagen and K-70 gels. Our assumption was
that directed fusion of closely-placed cell aggregates
should be observed with higher affectivity on the collagen
gel than the thermo-reversible gel. It was shown that
placing cell aggregates in close opposition on the surface of
collagen type 1 gel resulted in the fusion of adjacent cell
aggregates, with a sequential formation of elongated rodlike tissue constructs (see Fig. 4). Similar results were
produced by placing the cell aggregates on the surface of
collagen, and sequentially covering the aggregates with a
second collagen layer. This proved that fusion of cell aggregates occurs not only on the surface but also within 3D
collagen gel. Finally, we attempted to reproduce the same
results using the thermo-reversible gel. Although the
overall effectiveness of the thermo-reversible gels in promoting cell aggregate fusion was not equal to that of the
collagen type 1 gel, we were able to show that fusion does
occur in these gels.
To show that directed fusion can occur on patterned
gels, we loaded the collagen gel into the print head of the
printer. The gels were deposited in ring patterns about 1
cm in diameter and 500 ␮ wide. After gelling occurred, the
BAEC aggregates were added to the gel. Observation after
24 hr revealed that the cells were spread throughout the
gel. A live/dead assay image is shown in Figure 5, which
shows that the cells within the gel were alive, whereas
most of the cells that migrated out of the gel were not
Directed Cell Placing as the Main Problem of
Tissue Engineering
A very important task of tissue engineering is to put
certain cell types in exact places. This cannot be achieved
by traditional bioreactor-based cell-seeding technology using porous biodegradable scaffolds (even sophisticated
ones created with prototype printing). Other techniques,
such as rolling and embedding, are as yet not suitable for
engineering complex multicellular 3D organs.
Cell-printing technology using thermo-reversible gel
provides (at least theoretically) a unique opportunity to
solve this most difficult tissue-engineering task, and
opens the way for mass organ printing. This may eventually lead to a cost-effective solution to the problem of organ
Fig. 2. Top view of seven alternate layers of clear and tryan bluedyed K-70 gel on a collagen-coated dish. The edges between the layers
reveal the amount of mixing that occurs when a cold drop liquifies some
of the gel before it is gelled itself. Optimizing gelling kinetics, drop size,
and deposition rate may minimize this effect.
Fig. 3. Image of a single endothelial cell aggregate of ⬃670 ␮m
shortage. To assess the feasibility of this exciting technology, we attempted to remove some of the most obvious
technological barriers to its use.
Estimating the Optimum Minimal Thickness of
the Gel Layer
The presence of the gel layer between the cell aggregates serves two purposes: 1) it provides mechanical
strength and stability to the construct, and 2) it serves as
a drug-delivery device. For example, porous PLA/PGA/
PEG polymeric microspheres with controlled degradation
times could be put loaded into the printer for controlled
delivery of growth factors. Using copolymers of different
Fig. 4.
Cell aggregate fusion on the (A) collagen gel and (B) K-70 gel.
Fig. 5. Cell aggregate fusion on a printed collagen gel ring. The
image obtained on an epifluorescent microscope shows fused aggregates and cells that migrated across the gel. Cells that migrated toward
the center of the ring (in red) proved nonviable, possibly because they
became dry outside the gel.
PLA/PGA compositions, microspheres with a range of degradation times could be obtained. Porous microspheres
could be loaded with desired growth factors and dispersed
within an injectable gel matrix. The release profile of
bioactive agents could be adjusted to the required tissuespecific spatial/temporal pattern of expression.
Generation of Sequential Thermo-Reversible
Gel Layers
Apart from tissue-specific cell–matrix interactions, the
following aspects must be considered: gelation kinetics,
matrix resorption rate, possible toxicity of degradation
products and their elimination routes, and possible interference of the gel matrix with histogenesis.
There are several possible mechanisms that lead to in
situ gel formation: gelation in response to temperature
change (thermal gelation), pH change, ionic cross-linking,
or solvent exchange (Guenet, 1992). The kinetics of gelation are directly affected by the mechanism: thermal gelation, which is rate-limited by heat transfer, is faster
then ionic cross-linking and pH-triggered gelation, which
are rate-limited by mass transfer. The time it takes for the
gel to form will have an effect on cell spacing and distribution within the printed gel matrix. Gels that form in
response to temperature change may offer specific advantages related to fast gelation kinetics.
The thickness of the gel layers therefore depends on the
kinetics of the gel, as well as on the drop size, which is
governed by surface tension and needle diameter. To increase the efficiency of the technology, the gel thickness
will need to be adjusted to match the size of the aggregates. Using a needle with the same internal diameter as
the aggregates appears to give satisfactory results. However, as indicated in Figure 2, some distortions in the
layers (possibly due to a melting process) are observed.
Further improvements in controlling the aggregate size
distribution would also require a more rigorous model for
gel thickness. This model would need to take into account
gelation kinetics, surface tension, and interfacial forces
between the liquid and gel states of the solution.
Identification of the Optimum Size for Cell
The optimum size for cell aggregates is defined as a
diameter that is small enough (usually ⬍1 mm) to allow
centrally positioned aggregate cells to survive, and a magnitude that is equal to the thickness of the successively
printed gel layers. The second restraint is much more
important because it determines the overall feasibility of
the technique, whereas apoptosis can be prevented either
by using telomerized cells or cells transfected with blc-2
(Yang et al., 2001; Nor et al., 2001). Based on estimations
of the optimum minimal thickness, the diameter of the cell
aggregates must be no smaller than 600 ␮m. This does not
exclude the possibility of variability in cell aggregate size,
but it definitely places a well-defined limit on the smallest
technologically acceptable diameter.
Role of Gel and Close Positioning in Promoting
the Fusion of Cell Aggregates
The fusion of cell aggregates is a spontaneous phenomenon in aggregate suspension. Recently, important insights have been gained regarding the possible mechanism of aggregate fusion in the presence of extracellular
matrix (Ryan et al., 2001). It was demonstrated that competition exists between cell-to-cell forces vs. cell-to-substrate forces. If the cell-to-substrate force is stronger than
the cell-to-cell force, the outcome will be a monolayer. If
the force of cell-to-cell contacts is predominant, the cells
will form aggregates. This principle was used by us when
we generated cell aggregates. The absence of a cell adhesive substrate forced the cells to aggregate in order to
interact with each other.
Our data demonstrate that collagen gels are more aggregate-friendly and are more “fusogenic” substrates com-
pared to thermo-reversible gels. This is a more complex
situation than the single-cell or monolayer case described
above. It is possible that the tethering effect of collagen
mediated by the RGD tripeptide is responsible for keeping
cell aggregates together, and for their sequential fusion.
We are exploring ways to optimize the thermo-reversible
gel by incorporating RGD or other extracellular matrix
analogs (Mann et al., 2001; Griffith and Naughton, 2002)
without compromising its thermo-reversible properties.
Toward Organ Printing
Keeping in mind that the adhesiveness of gels for cells
can be modified, our data show that cell aggregate fusion
can be achieved in the thermo-reversible gel. Our data
also demonstrate that we can reproducibly generate sequential layers of thermo-reversible gels of a thickness
corresponding to the cell aggregate diameter. This thickness allows cell aggregates to be laced in closely enough to
fuse. Both 3D collagen and thermo-reversible gels were
shown to allow fusion of closely-placed cell aggregates.
Aggregate fusion causes the 3D structures to shrink somewhat, an effect that will need to be considered in the
design of organ blueprints. However, we can also assume
that additional extracellular matrix will be formed by the
cells, thus providing additional glue for cell aggregate
fusion and serving as a sort of tethering device for promoting the self-assembly of printed tissue constructs. Furthermore, this additional extracellular matrix may act as
a buffer between adjacent structures, and thus be more
forgiving with respect to the exact tolerances of a blueprint.
In addition, the use of gel with RGD ligands may enhance adhesion. Alternatively, aggregates trapped inside
gel drops can be printed. Because closely placed drops
fuse, we expect the aggregates to fuse as well. However,
the time frame of cell attachment is very important, since
it determines the speed at which successive layers can be
printed. Soft-tissue adhesives such as cyanoacrylate esters, fibrin sealant, and gelatin-resorcinol-formaldehyde
glues, or other bioadhesives, could dramatically prevent
the constructs from being washed off during successive
printing cycles. The addition of a corresponding growth
factor in slow-release microspheres could accelerate and
direct this process. The programmed release of different
growth factors in various concentrations and sequences is
also technically possible. Finally, the thermo-reversible
gel can be optimized or functionalized by the addition of
extracellular matrix analogs. This would promote cell adhesion, and result in better cell survival as well as more
effective cell aggregate fusion.
Poor cell survival in the printed aggregates is a possible
drawback to the described technology. We are currently
exploring different ways to avoid apoptosis and necrosis
inside the aggregates, such as adding survival factors
(e.g., basic fibroblast growth factor), using transient genetic modifications of cells with antiapoptotic (e.g., bcl-2
and telomerase), and blocking apoptotic pathways. It is
well known that the survival of cells in cell aggregates can
be seen as a result of cell contact-mediated survival stimuli. In this respect, we expect the survival rate of aggregated cells to be higher than that of single printed cells.
Furthermore, this technology can tolerate a small percentage of cell death.
The goals of this study were to demonstrate the feasibility of organ-printing technology, solve some important
technical problems, and eliminate some critical technological barriers. The actual printing of 3D tissue constructs
will be the subject of another paper, in which we will also
address issues of cell survival, tissue perfusion, and vascularization.
Taken together, our data strongly indicate that the proposed organ-printing technology is feasible using the originally-designed cell printer and thermo-reversible gel.
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