Rapid prototyping in tissue engineering: challenges and potential

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MMII is selected by researchers to fabricate calcium
phosphate scaffolds because the building material has a
very low coefficient of thermal expansion. There will be
minimal risk of fracture due to coefficient of thermal
expansion mismatch with the ceramic during pyrolysis.
Limpanuphap and Derby [52] fabricated TCP scaffolds
with controlled internal porosity using a suspension of
TCP in an acrylate binder. A similar route was used to
produce a polymer–TCP scaffold, which is believed to show
more potential for cell adhesion. Wilson et al. [53]
fabricated HA scaffolds with a defined macroarchitecture.
Channels achieved are w350–400 mm wide. These authors
suggested that the inherent surface texture obtained from
an MMII-built mould increased the surface area of the
scaffold and led to a higher degree of calcium and phosphate
release, which is advantageous to bone formation.
Sachlos et al. [54] have successfully produced collagen
scaffolds with predefined and reproducible internal channels.
The smallest channel width achieved was reported to
be as low as 135 mm. In a variation to Sachlos’ work, the
authors have produced chitosan-collagen scaffolds. MMII
moulds with intricate channels were built to contain a
chitosan–collagen solution. The resultant gel-like scaffold
was capable of attaining the designed morphology
(Figure 4). These channels can serve as flow channels
when coupled with a customized bioreactor, achieving
more efficient perfusion of the culture medium.
Photopolymerization techniques
For photopolymerization techniques, optical energy is
applied to irradiate the thin layer at the surface of a
liquid photopolymer resin. The irradiation areas of
resin react chemically and transform into a solid
phase. A well-known photopolymerization technique is
stereolithography (SLA) [20].
SLA: A UV laser traces out the first layer of photocurable
resin, solidifying the model’s cross-section while
leaving the remaining areas in liquid form. The elevator
then drops by a sufficient amount to cover the solid
polymer with another layer of liquid resin. A sweeper
recoats the solidified layer with liquid resin and the laser
traces the second layer atop the first.
Chu et al. [50] have produced HA-based porous implants
using SLA-built epoxy moulds. A thermal curable HA–
acrylate suspension was cast into the mould to obtain a
scaffold with interconnected channels. The resolution of
channel width achieved was as low as 366 mm.
In another study, investigators carried out an in vivo
study using two different architecture designs, orthogonal
Figure 4. Scaffold (right) produced using ModelMakerIIe-built mould (left). The
scaffold is able to reproduce the predetermined morphology of the mould. The
channels ensure high diffusion efficiency across the scaffold.
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Review TRENDS in Biotechnology Vol.22 No.12 December 2004
and radial channels [55]. The preliminary results showed
that controlling the overall geometry of the regenerated
bone tissue was possible through the internal architectural
design of the scaffolds.
Challenges of RP in tissue engineering
In spite of the increasing interest of tissue engineers in the
use of RP, there are several challenges that need to be
addressed, namely the limited range of materials, the
optimal scaffold design, the bioactivity of the scaffold, as
well as the issues of cell seeding and vascularization. Each
of the issues will be discussed in detail.
Range of materials
Material processability: RP techniques are very
specialized technologies in terms of material processability.
Each technique requires a specific form of input
material such as filament, powder, solid pellet or solution.
Therefore, it must be ensured that the choice of materials
for the scaffold is compatible with the selected RP process
and that it can be efficiently produced in the form
required. Other considerations during the selection of
materials include the degradation profile and the mechanical
strength of the scaffold.
Degradation rate: In an ideal case, the scaffold should
be remodeled and resorbed by growing cells and gradually
replaced by the newly formed extracellular matrix and
differentiated cells. A desirable feature would be synchronization
of the polymer degradation rate with the rate of
tissue ingrowth. Therefore, the degradation properties of a
scaffold are of crucial importance for the success of the
scaffold-based approach.
The degradation–absorption mechanism is the result of
many interrelated factors, including the hydrophilicity of
the polymer backbone, degree of crystallinity, presence
of catalysts, volume of porosity and the surface area.
Balancing each of these factors will enable an implant to
degrade slowly and transfer stress at an appropriate rate
to the surrounding tissues as they heal. This is one of the
major challenges facing tissue engineering research today.
Degradation product: Even though degradation
products of biodegradable polymers are known to be
largely non-cytotoxic, little information is available
regarding the degradation rate-dependent acidic byproduct
effect of the scaffold. Sunga et al. [56] found that fast
degradation of the polymer negatively affects cell viability
and migration into the scaffold, both in vitro and in vivo.
This can be explained by the rapid local acidification due
to polymer degradation. Therefore, a more systematic
investigative approach is needed to classify the material
degradation profile.
Mechanical strength of scaffolds: Cells are able to
detect with high sensitivity the mechanical properties of
the adhesion substrate, and to regulate integrin binding
and the assembly of focal adhesion plaques and the
cytoskeleton accordingly [57]. If the adhesion substrate
is too rigid and nondeformable, the cells are not able to
reorganize and recruit the receptors into focal adhesion
plaques, which is a prerequisite for the delivery of signals,
ensuring the viability of anchorage-dependent cells.
Similarly, if the material is too compliant, it does not