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Tissue engineering (TE) aims at reproducing morphogenesis in the laboratory, i.e., in vitro, to fabricate replacement organs for regenerative medicine. The classical approach to generate tissues/organs is by seeding and expanding cells in appropriately shaped biocompatible scaffolds, in the hope that the maturation process will result in the desired structure. The novel technology relies on the concept of tissue liquidity according to which multicellular aggregates composed of adhesive and motile cells behave in analogy with liquids. To emphasize the major role played by tissue fusion in the embryo and explain how the parameters (surface tension, viscosity) that govern tissue fusion can be used both experimentally and theoretically to control and simulate the self-assembly of cellular spheroids into 3D living structures. The experimentally observed postprinting shape evolution of tube- and sheet-like constructs is presented. Computer simulations, based on a liquid model, support the idea that tissue liquidity may provide a mechanism for in vitro organ building. As for scaffolding TE, there were studies done on mouldless manufacturing techniques, which are known as solid free-form fabrication (SFF), or rapid prototyping, have been successfully used to fabricate complex scaffolds. Similarly, to achieve simultaneous addition of cells during the scaffold fabrication, novel robotic assembly and automated 3D cell encapsulation techniques are being developed. As a result of these technologies, tissue-engineered constructs can be prepared that contain a controlled spatial distribution of cells and growth factors, as well as engineered gradients of scaffold materials with a predicted microstructure. Therefore by comparing both of these TE methods we are able to know that which method is the best for future TE on pharmaceutical usage.
Tissue engineering is a type of regenerative medicine where it is working its way to restore structure and function of damaged tissues and organs. It is also working to create solutions for organs that become permanently damaged (Wikipedia, 2009a).
Scientific research is working to make treatments available for clinical use. Treatments include both in vivo and in vitro procedures. In vivo meaning studies and trials performed inside the living body in order to stimulate previously irreparable organs to heal themselves. In vitro treatments are applied to the body through implantation of a therapy studied inside the laboratory (Wikipedia, 2009a).
Definition of Tissue Engineering.
There are 2 different types of methods for this in vitro tissue engineering. Which are Bio-printing and Scaffolding technique. In bio printing or in other words, it is called as organ printing, it is a biomedically relevant variant of rapid prototyping technology, which is based on tissue fluidity. Where it is computer-assisted deposition (it is called 'printing' in short) of natural materials (cells or matrix) is done one layer at a time until a particular 3D form is achieved. Whereas for scaffolding, cells that are implanted or 'seeded' into an artificial structure capable of supporting three-dimensional tissue formation. The most comment cells used for tissue engineering are stem cells (Wikipedia, 2009a).
Stem Cell Defination.
Stem cells are cells found in most multi-cellular organisms. It has the ability to choose between prolonged self-renewal and differentiation. This fate choice is highly regulated by intrinsic signals and the external microenvironment, the elements of which are being rapidly elucidated (Wikipedia, 2009b).
Stem cells can be found in many adult mammalian tissues. For example, epithelia, blood, and germline, stem cells contribute to replenishment of cells lost through normal cellular senescence or injury. Stem cells may also be present in other adult organs, such as the brain and pancreas, which normally undergo very limited cellular regeneration or turnover. Stem cells in adult tissues may have more "plasticity" than originally thought, they typically form only a limited number of cell types. In contrast, stem cells of the early mammalian embryo, have the potential to form any cell type. In the unmanipulated blastocyst-stage embryo, stem cells of the inner cell mass (ICM) promptly differentiate to generate primitive ectoderm, which ultimately differentiates during gastrulation into the three embryonic germ (EG) layers. ICM cells are the source cells from which pluripotent mouse, nonhuman primate, and human embryonic stem (ES) cells are generally derived, although there is evidence that mouse ES cells may be more closely related to primitive ectoderm (Wikipedia, 2009b).
Types of stem cells.
Stem cells are divided in to 2 types which are embryonic and adult stem cells. Whereby in embryonic stem cells are derived from then embryo as the name suggested. Embryonic stem cells are derived from embryos that develop from eggs that have been fertilized in vitro environment. Thus they are not derived from the egg fertilized in a subject or sample. It goes the for stem cells that has been derived from the human, the human embryonic stem cells are typically four or five days old and are a hollow microscopic ball of cells called the blastocyst (Wikipedia, 2009b).
Blastocyst is a structure formed in the early embryogenesis of mammals, after the formation of the morula, but before implantation. The blastocyst includes three structures: the trophoblast, which is the layer of cells that surrounds the blastocoel, a hollow cavity inside the blastocyst; and the inner cell mass, which is a group of cells at one end of the blastocoel that develop into the embryo proper (Wikipedia, 2009b).
Adult stem cells are cells that have not yet developed into specialized cells. It is found among differentiated cells in a tissue or organ that can renew itself and can differentiate to form some or all of the major specialized cell types of the tissue or organ. The main functions of these adult stem cells are to maintain and repair the tissue in a living organism. Adult stem cell can be found in many places of a living organism which include, brain, bone marrow, peripheral blood, blood vessels, skeletal muscle, skin, teeth, heart, gut, liver, ovarian epithelium, and testis (Wikipedia, 2009b).
Differentiation Between Embryo Stem Cell and Adult Stem Cell.
Embryonic and adult stem cells both have their own advantages and disadvantages. One major difference between adult and embryonic stem cells is their different abilities in the number and type of differentiated cell types they can become. Embryonic stem cells can differentiate into all types of cell in the living organism. Whereas Adult stem cells are limited to differentiate into different cell types of their tissue of origin (Stem cell, 2009).
Embryonic stem cells can be grown relatively easily in culture. Adult stem cells are rare in mature tissues, so isolating these cells from an adult tissue is challenging, and methods to expand their numbers in cell culture have not yet been worked out. This is an important distinction, as large numbers of cells are needed for stem cell replacement therapies (Stem cell, 2009).
Introduction of Bio-Printing.
Bio-printing is defined as a rapid prototyping computer-aided 3D printing technology, whereby the deposition of cell or cell aggregate where deposited to a layer by layer conformation into a 3D gel with sequential maturation of the printed construct into perfused and vascularized living tissue or organ. This bio-printing includes many different printer designs and components of the deposition process such as, jet-based cell printers, cell dispensors or bioplotters, the different types of 3D hydrogels and varying cell types (Marga et al., 2007).
The cell aggregates are actually cells or more specific stem cells that act as liquid properties, whereby these liquid are the bio-ink for this bio-printing. The hydrogel was the bio-paper. It is made out agarose, neurogel and collagen. Next, the self-assembly process was left to occur, where by the formation of a 3D organ begin to take place (Marga et al., 2007).
In the process of 3D bio-printing, spatial patterning of hormones, including growth factors is known to be critical in directing all aspects of cell behavior throughout life, including embryogenesis and wound repair. Simple, controllable methods to engineer the physical placement and concentration of immobilized exogenous growth factors in a physiologically relevant manner and to study cell behavior in register to such persistent patterns are important for biological research, as well as being a logical consideration for developing tissue regeneration therapies. Therefore, we have developed an inkjet-based bio-printing system and methodologies are developed to create concentration-modulated two-dimensional patterns of growth factors immobilized onto fibrin substrates using native binding affinities (Marga et al., 2007).
Experiment on Bio-printing Techniques.
Spherical cell aggregates act as the shape of a liquid droplet adopts on nonadhesive substrate under no external forces. Therefore Confluent Chinese hamster ovary (CHO) cell cultures transfected stably with N-cadherin were grown in Dulbecco's Modified Eagle Medium (DMEM) supplemented with 10% Fetal Bovine Serum, 10 µg/ml penicillin, streptomycin, gentamicin, and kanamycin, 400 µg/ml geneticin, were washed twice with Hanks' balanced salt solution (HBSS) containing 2 mM CaCl2, then treated for 10 min with trypsin 0.1% . Depleted cells were centrifuged at 2,500 rpm for 4 min (Jakab et al., 2004).
The pellet was then transferred into capillary micropipettes of 500 µm diameter and incubated at 37°C with 5% CO2 for 10 min until they recovered their adhesive connections. The firm cylinders of cells removed from the pipettes were cut into 500 µm long fragments then incubated in 10 ml tissue culture flasks with 3 ml of DMEM on a gyratory shaker at 120 rpm with 5% CO2 at 37°C for 24-36 h. This procedure reproducibly provides spherical aggregates of similar size (500 µm diameter) (Jakab et al., 2004).
Biocompatible gels can be used either to promote natural regeneration by cell recruitment in vivo, or to serve as a supporting environment in creation of an artificial implantable tissue in vitro. In the first case the scaffold provides support in vivo; in the second one, this support is provided until the cells are assembled and maturated enough to support themselves. Chemical properties of the scaffold pore size, biodegradability, cell-gel adhesion molecules, growth factors, immunoresponse and cytotoxicity are factors to consider in biocompatible gel design (Jakab et al., 2004). The materials were:
Neurogel is a biocompatible porous poly [N-(2-hydroxypropyl) methacrylamide] hydrogel. This gel has been shown to provide optimal conditions for spinal cord repair, it contains Arg-Gly-Asp (RDG) fragments, a protein with affinity to integrins, a special class of cell-matrix adhesion molecules (Jakab et al., 2004).
Three Dimensional Printer.
The printer needed in this bio-printing project is Roland MDX-20. Where it is a 3D printer which is capable in printing out any object or prototype, base on the type of the ink that is used by this printer. For example, polymer which can get harden can be shaped in to any prototypes like tree. Therefore, the 3D formation of the cell aggregates is possible (Jakab et al., 2004).
Immobilization of Growth Factor (Bone Morphogenetic Cell) to the Muscle Derived Stem Cells.
Preparation of Fibrin-Coated Glass Slide.
A glass slide was cut into squares shaped. Then these squares were cleaned in a sulfuric acid bath with NOCHROMIX for 2 h, rinsed 10 times with deionized water, and dried under nitrogen gas. Cleaned squares were then incubated in a 95% acetone solution containing 1% 3-aminopropyltriethoxysilane for 10 min at 23°C to functionalize the glass surfaces with amine groups. Subsequently, squares were rinsed 3 times respectively in acetone, ethanol, and deionized water to hydrolyze the silane and quench the surface reaction (Qu-Petersen et al., 2004).
Silanized squares were dried at 120°C for 45 min and incubated in a 3% glutaraldehyde solution in PBS, pH 7.4, for 2 hour at 37°C to react with the amines on the glass surfaces and expose reactive aldehyde groups. Squares were then rinsed twice with methanol and deionized water, respectively (Qu-Petersen et al., 2004).
Glutaraldehyde prepared squares were coated with fibrinogen by incubation in 0.1 mg/ml fibrinogen contained in 10 mM sodium phosphatebuffer, pH 7.4 for 18 h at 4°C. Excess unbound fibrinogen was removed by aspiration and the remaining active sites blocked with 0.3 M glycine, pH 7.4 for 2 hour at 4°C followed by three rinses with PBS. Immobilized fibrinogen was converted into fibrin by incubating squares in 4 U/ml thrombin contained in 10 mM sodium phosphate buffer with 1mM calcium chloride at 37°C for 2 h, then rinsed 3 times with PBS and stored in PBS at 4°C for up to 2 weeks. Prior to patterning, squares were rinsed 3 times with sterile deionized water and air-dried in a laminar flow hood. The presence and uniformity of the fibrin coating was verified using transmission and scanning electron microscopy. The thickness of the fibrin films was estimated to be approximately 20 nm. (Qu-Petersen et al., 2004)
Bio-Ink and Bio-Paper Preparation from BMP-2.
The preparation of bio-ink had been shown in the previous step. Instate of using Chinese hamster ovary cells, Bone morphogenetic protein 2(BMP-2) has used. Then the BMP2 was labeled with cyanine 3 which is a monofunctional reactive dye. Cy-3-labeled BMP-2 was purified from free dye over a heparin-Sepharose column and stored at -80°C to maintain viability. In order to sterilize the printer head, ethanol was used to rinse and then followed by 3 rinses of 0.2 µm of filtered deionized water. The bio-ink, which consisting of 100 µg/ml Cy3-labeled BMP-2 in phosphatebuffered saline (PBS), pH 7.4, was loaded into the printhead, and patterns were printed onto fibrin-coated glass slide (Phillippi et al., 2007).
These relatively high growth factor bio-ink concentrations were used to aid visualization of the patterns. The printed patterns were 2 ? 2 arrays of 750-µm squares with a spacing of 75 µm between the printed drops and 1.75 mm between the squares. Squares were printed as 2, 8, 14, and 20 overprints of the bio-ink to modulate the surface concentrations (Phillippi et al., 2007).
The patterns were then incubated at 37°C with 5% CO2 and 95% humidity in serum-free Dulbecco's modified Eagle's medium (DMEM) with 1% penicillin/streptomycin and were imaged every 24 h up to 144 h by epifluorescence microscopy with a Cy3 filter set (Phillippi et al., 2007).
Muscle Derived Stem Cells (MDSCs) were isolated as previously described and cultured in DMEM (high-glucose), 10% fetal bovine serum (FBS), 10% horse serum, 1% chick embryo extract, 200 mM glutamine, and 1% penicillin/ streptomycin including prophylactic for myoplasma. C2C12 were obtained from American Type Culture Collection (Manassas, VA, http://www.atcc. org) and cultured according to the manufacturer's instructions (DMEM [high-glucose], 10% FBS, 1% penicillin/streptomycin). Where noted, Noggin (100 ng/ml as conditioned media) or 50 ng/ml FGF-2 was added to the cultured media. Doses of FGF-2 and Noggin were determined on the basis of prior studies. For myogenic conditions, serum levels were adjusted to 1% FBS and 1% horse serum (Celil et al., 2005).
The MDSCs are cells isolated from a adult mice, where the organs are fully developed which means the stem cells have already developed into a specialize cells which are capable in performing some task (Celil et al., 2005).
As for C2C12 cells they are embryonic stem cells which are isolated from the mice where it can be purchased from the American Type Culture Collection as mention above. This C2C12 cells are the basic cell units where it has not transformed into specialize cells as it had been mention previously (Celil et al., 2005).
Gene Expression of alp and osx Gene by Quantitative Real-Time Polymerase Chain Reaction.
The optimum concentration of the BMP-2 for the immobilization of the muscle derives stem cells (MDSCs) is unknown. Therefore the gene expression test is done in order to get its optimum concentration for the immobilization of the MDSCs to work efficiently (Jadlowiec et al., 2004).
The concentration of BMP-2 was prepared with 125, 250, or 500 ng in 10 µl H2O was then hand-pipetted onto collagen-coated dishes (12 wells) and allowed to air-dry in a sterile hood. Plates with blotted BMP-2 were then incubated overnight in complete basal medium to release unbound BMP-2. Cells were then seeded to 90% in basal medium on the BMP-2 blots. BMP-2 was added to the culture media in parallel wells (125, 250, and 500 ng/ml) as positive control. For negative control, cells were cultured in basal medium alone (no BMP-2). Cells were cultured for 24 h prior to extraction of total RNA (Jadlowiec et al., 2004).
Quantitative Real Time PCR protocols are performed in order to observe the gene expression of Alp and Osx gene (Figure 3). Where by the 1st step is extraction of the RNA from the cells, in this case BMP-2. Then followed by the selection of dye, where Taqman probe are selected. 10-30 ng of total RNA were added per 50-µl reaction with sequence-specific primers (200 nM) and Taqman® probes (200 nM). qPCR assays were carried out in triplicate on an ABI Prism 7000 sequence detection system (Jadlowiec et al., 2004).
Then the process of reverse transcription takes place at 40°C for 30 min. Followed by the process of initial denaturation at 95°C for 10 min. Then 40 cycle of denaturation at the temperature of 95°C for 15s, followed by annealing and extension process at 60°C for 45s ( Jadlowiec et al., 2004).
Prior to printing BMP-2, Alp and Osx gene expression were quantified in cells cultured on solid-phase blots of BMP-2 dried on collagen-coated tissue culture wells. Both C2C12 cells and MDSCs expressed basal levels of Alp gene expression that were dose-dependently upregulated in response to a 24-hour exposure of increasing concentration of BMP-2 dried and immobilized to collagen-coated tissue culture wells. The lowest dose of solid-phase BMP-2 (125 ng/10 µl deposited) does not show a significant increase in Alp gene expression relative to negative control for either MDSC or C2C12. However, in 250 ng/10 µl deposited stimulated significant upregulation of Alp in MDSC of 33.17-fold, but 500 ng/10 µl does not show a significantly different from 250 ng/10 µl deposited, with an upregulation of 26.51-fold. Both 250 and 500 ng/10 µl deposited were higher compared with negative control. Similarly, Alp upregulation was quantified in C2C12 as 3.18-fold for 250 ng of solid-phase BMP-2 and 17.8-fold for 500 ng/10 µl deposited relative to control. Neither MDSC nor C2C12 expressed basal Osx in this experiment. However, the highest dose of solid-phase BMP-2 (500 ng/10 µl deposited) was sufficient to induce detectable Osx gene expression in both MDSC and C2C12 (Jadlowiec et al., 2004).
Alkaline Phosphatase Activity.
Cells were seeded to more or less 90% on BMP-2 patterns. BMP-2 (50 ng/ml) was delivered to the media in parallel wells as the positive control. Basal medium was the negative control. Cells were cultured for 72 h and then rinsed in PBS and fixed for 2 minutes in 3.7% formaldehyde. Alkaline phosphatase (ALP) activity (SIGMAFAST) was detected according to the manufacturer's instructions (Phillippi et al., 2007).
Immunocytochemistry for Myosin Heavy Chain.
Cells cultured on BMP-2 patterns were fixed in 3.7% formaldehyde and stained immunocytochemically for myosin heavy chain (fast) (MHC-f) to detect the formation of myotubes. Cells were counterstained with 4,6-diamidino-2-phenylindole (Xu et al., 2005).
The formation of myotube is from the fusion of myoblasts (muscle stem cells) into multi-nucleated fibers. In the early development of an embryo, these myoblasts will proliferate if enough fibroblast growth factor is present. When the FGF runs out, the myoblasts cease division and secrete fibronectin onto their extracellular matrix (Xu et al., 2005).
MDSC and C2C12 have the capacity to differentiate toward the osteogenic lineage, evidenced here by induction of ALP activity by liquid-phase BMP-2 for 72 hours (Figure 4, A and B). MDSCs and C2C12 cultured on patterns of printed BMP-2 also differentiate toward the osteogenic lineage, as an evidenced the ALP activity had increased (Figure 4, A and B, respectively) in register to the printed BMP-2 patterns. A patterned differentiative response was visible by eye within tissue culture wells, as shown by the scanned images of the culture wells (Figure 4, A and B, middle). An increased magnification of the tissue culture wells shows the dose-dependent pattern response to cells on BMP-2 patterns for both MDSC and C2C12 (Figure 4, A and B, far right) (Xu et al., 2005).
MDSC and C2C12 osteogenic differentiation were dose-dependent with respect to the surface concentration of printed BMP-2, with the highest surface concentration inducing a near-perfect square region of differentiating cells that corresponds to the BMP-2 array (Figure 4, C and D, respectively), whereas cells seeded outside the spatially defined BMP-2 pattern did not exhibit appreciable ALP activity and thus presumably remain undifferentiated. Therefore the data provided proved that the concept for engineering spatially controlled engineered cell fate toward the osteogenic lineage may happen (Xu et al., 2005).
Osteogenesis is the process of laying down new bone material by osteoblasts. It occurs in two different ways. The 1st way is by Intramembranous Osteogenesis (see intramembranous ossification), which is the direct laying down of bone into the primitive connective tissue (mesenchyme), or the 2nd way is by Endochondral Osteogenesis (see endochondral ossification) which involves a cartilage precursor (Xu et al., 2005).
Next is to find out whether BMP signaling antagonists such as FGF-2 or Noggin present in the liquid-phase would influence cell response to the BMP-2 printed patterns shown is (Figure 5). Scientist found out that Noggin delivered to the media inhibited BMP-2 pattern response in both MDSCs and C2C12 cells (Xu et al., 2005).
The patterned cell differentiation in MDSCs were partially inhibited by FGF-2, whereas a complete inhibition in C2C12, where no appreciable ALP activity was detected in the presence of FGF-2 Figure 5A. Noggin present in the culture media likely binds to the printed immobilized BMP-2 and thus inhibited BMP-2 signaling on patterns. The binding of Noggin-BMP-2 binding present is due to cell attachment and BMP receptor activation, since Noggin was administered to the media at the time of cell seeding. In contrast, FGF-2 has been shown to elicit cross-signaling mechanisms with BMP (Xu et al., 2005).
Time-Lapse Image Acquisition.
The microscope stage incubator was mounted on a Zeiss IM35 Axiovert microscope using a 5X, 0.15 N.A. phase 1 objective and phase optics with a computer operated stage. A ground glass diffuser was used to smooth out the bright field illumination. Time-lapse imaging of the patterns was automated using BDS-Image (Biological Detection Systems) software and custom written software that runs within the BDS-image environment. The patterns (located using the marks scored on the slide after printing) were imaged every 30 min using a Photometrics C-250 cooled CCD camera, and the images were immediately processed by the software before the next image acquisition cycle. A total of 5 fields were imaged. Four of the fields contained the printed patterns, while the fifth field, adjacent to the 22 and 32 overprint squares, did not contain a printed pattern and served as a control. For the endpoint experiments where arrays consisting of 9 squares were printed, the cell response was imaged only once at 100 h (Qu-Petersen et al., 2004.).
The researchers took advantage of the inherent myogenic potential of the MDSCs to spatially control the differentiation of two lineages simultaneously in the same culture well using the inkjet printing approach. MDSCs that were cultured on 2?2 patterns of arrays of four different surface concentrations of BMP-2, which has been mention previously were then under myogenic culture medium conditions. MDSCs exhibited ALP activity in register to the printed BMP-2 patterns that has been shown in Figure 6 as before in normal serum in Figure 4. All printed patterns showed dose-dependent ALP response and no myotubes on pattern (Qu-Petersen et al., 2004.).
Engineering Biological Structures of Prescribed Shape using Self Assembling Multicellular Systems.
In tissue engineering, the studies of cell assembly are essential, where by the structure of the tissue or even organ can be formed from biofabrication. Self-assembly is a fundamental process that drives structural organization in both inanimate and living systems. It is in the course of self-assembly of cells and tissues in early development that the organism and its parts eventually acquire their final shape (Whitesides and Grzybowski, 2002.).
Histogenesis and organogenesis are examples of self-assembly processes, in which, through cell- cell and cell- extracellular matrix interactions, the developing organism and its parts gradually acquire their final shape (Whitesides and Grzybowski, 2002.).
Histogenesis is the developmental processes by which the definite cells and tissues which make up the body of an organism arise from embryonic cells. Among animals, the ectoderm, endoderm, and mesoderm, also known as the primary germ layers, provide the stem cells which gradually transform into distinctive kinds of cells and tissues. In the higher plants, meristematic cells, which occur wherever extensive growth takes place, provide the basis for tissue formation (Whitesides and Grzybowski, 2002.).
Organogenesis is the process by which the ectoderm, endoderm, and mesoderm develop into the internal organs of the organism. Internal organs initiate development in humans within the 3rd to 8th weeks in uterus. The germ layers in organogenesis differ by three processes: folds, splits, and condensation. Developing early during this stage in chordate animals are the neural tube and notochord. For instant vertebrate animals all differentiate from the gastrula the same way. Vertebrates develop a neural crest that differentiates into many structures, including some bones, muscles, and components of the peripheral nervous system. The coelom of the body forms from a split of the mesoderm along the somite axis. (Whitesides and Grzybowski, 2002.)
Cell Aggregate Preparation.
The preparation of cell aggregate is the protocol as the preparation of tissue liquid which is the bio-ink. Brightfield and phase contrast microscopies were not suitable for imaging the fusion of embedded aggregates due to the opaqueness of some gels, thus all aggregates were made from cells transfected with histone attached yellow fluorescent protein. To visualize interpenetration of cells between adjacent aggregates and 3D tubular structures, the aggregates were stained with PKH2 and PKH26 (green and red) membrane intercalating dyes. These fluorescent markers do not influence the adhesive properties of cells, as revealed by cell sorting studies. (Foty et al., 1996)
From the previous experiment, the preparation of bio-paper need to undergo a series of selection before it is been used as a bio-paper. This is to ensure that the bio-ink which are the cell aggregates are bio compatible to the medium which will be printed on (Jakab et al., 2004). The few medium includes:
Agarose. 4% agarose solution was prepared by dissolving UltraPure Low Melting Point Agarose in distilled water. At this concentration the agarose solution gels at 28-30°C, excluding heat damage of the cells during embedding.
NeuroGel. Neurogel is a biocompatible porous poly [N-(2-hydroxypropyl) methacrylamide] hydrogel. This gel has been shown to provide optimal conditions for spinal cord repair it contains Arg-Gly-Asp (RDG) fragments, a protein with affinity to integrins, a special class of cell-matrix adhesion molecules.
Collagen. Rat tail collagen Type I was dissolved in 1 M acetic acid. Ham's F12 medium was added to promote cell viability, and then pH of the solution was adjusted to 7.0 with sodium bicarbonate and HEPES (4-(2-hydroxyethyl)-1-piperazineethanesulfonic acid) buffer solution. Depending on composition, this mixture of gels is left at room temperature in a few minutes. To tune the cell-gel interaction, three concentrations were prepared, at 1.0, 1.2 and 1.7 mg/ml final concentrations.
From these materials, an experiment has done to choose the most compatible materials for the production of bio-paper. From the preparation above, NeuroGel disks of 10 mm diameter and 2 mm height were washed 3 times in DMEM to eliminate the storage medium. A 0.5 mm wide, 0.5 mm deep circular groove was cut into a disk, then filled with 10 contiguously placed aggregates. The groove was then refilled with gel to completely embed the aggregates. The structure was kept in incubator for 3 days at 37 °C and 5% CO2 in a tissue culture dish containing 10 ml of DMEM (Jakab et al., 2004).
The gel-aggregate structures in agarose and collagen gels were created by placing a continuous ring of 10 aggregates or rectangular sheets of 25 aggregates on the top of a previously solidified gel layer, then covering with liquid gel that embedded the aggregates after gelation. The samples were incubated under the same conditions as described above (Jakab et al., 2004).
Then the cells aggregate were monitored by using an imaging technique. The Aggregate structures in transparent agarose and collagen gels were visualized on an Olympus IX-70 microscope with fluorescent attachment at 4X magnification, images were captured with a Nikon CoolPix 5000 digital camera. Phase contrast microscopy images were taken on the constructs embedded in collagen gels (Jakab et al., 2004).
Structures prepared in NeuroGel were washed in PBS solution then embedded in Tissue-Tek OCT Compound. The samples were slowly cooled (1°C/min) to -20°C in Nalgene freezing container, then cryosectioned on a yielding 10-16 slices mounted on microscope slides. Slices were visualized by fluorescent microscopy (Jakab et al., 2004).
The ability of aggregates to fuse depends on the mutual properties of the cells and the gel. Agarose represents a highly nonpermissive scaffold, cells do not interact at all with the environment, being unable to migrate into or reorganize the gel, and thus the initial and final configurations do not differ, the structure remains "frozen" in the initial state. Collagen 1.0 and 1.2 mg/ml as well as RGD containing NeuroGel match more the definition of the nonpermissive gel. These gels favor much less (collagen) or not at all (NeuroGel), the dispersion of cells into the scaffold, thus facilitating fusion. A high concentration of collagen is analogous to a permissive environment, the cell-gel interactions dominate over the cell-cell ones, the construct tends to collapse into a single clump of cells, in striking analogy with the rounding-up behavior of irregular tissue fragments placed into a non adhesive environment. Therefore, gels which contain collagen are much biocompatible to the cell aggregate, thus collagen gel had been chosen as the most suitable medium (Jakab et al., 2004).
The Influence of the Scaffold on 3D Structure Formation.
The above results demonstrate that scaffold properties affect cellular structures. The specific mechanism of how the gel influences pattern formation depends on its detailed chemistry. From the journal that had been published by Harris, A. K. and his co-workers it says that cells exert traction forces on their substrates and surrounding 3D matrices. Therefore, collagen has been a subject of numerous studies on this phenomenon (Vernon and Sage, 1996).
The pattern formation process in this case has some interesting features. As the fusion of the aggregates takes place, the ring of aggregates noticeably contracts, at least until approximately 60 h as it has shown in Figure 10 (Fray et al., 1998).
For the higher concentrations of collagen the contraction is more dramatic, whereas for the nonpermissive agarose and NeuroGel no similar effects were observed. The contraction results from the Chinese Hamster Ovary (CHO) cells pulling on the collagen. At some point in the fusion process the pattern assumes a starburst appearance, noticeable at 60 h but not before Figure 11. The spikes or bundles of cells extend outside of the ring, suggesting that by this time a network of radially aligned collagen fibers has developed in the vicinity of the aggregates (Fray et al., 1998).
In contrast, no similar effect is observed inside of the ring, not even at 144 h Figure 11. The probable reason collagen fibers do not align inside the ring is that the vectorial sum of the isotropically acting traction forces is zero. An analogous cancellation makes the electric field in the interior of a conducting spherical shell to be zero (Fray et al., 1998).
The contraction was quantified by measuring the area enclosed by the outer perimeter of the fluorescent ring. Curves represent exponential fits to the data of the form A exp(-t/tCG)+B, A and B being positive constants. The characteristic timescale of the contraction is defined by tCG, approximately 57 hours for the 1.0 mg/ml collagen concentration, used in our analysis (Fray et al., 1998).
The transparency of the collagen can be viewed by using brightfield microscopy to detect the aggregate fusion. That is defined as a measure of fusion the instantaneous value of the angle formed by two aggregates. Figure 11 shows the variation of the boundary between two adjacent aggregates for the 1.0 mg/ml collagen gel. Initially the angle between the tangents drawn to the joining point is zero, approaching 180 degrees when the fusion is completed. The curve is an exponential fit of the form C[1-exp(-t/tCC)], with C positive constant and tCC = 23 hours, the characteristic timescale of the aggregate fusion (Fray et al., 1998).
Optimization of Fusion by the Geometry of Initial Configuration: Sheet Formation.
In the previous experiment the structure formation was placed in a circular pattern. This initial geometry does not affect the time evolution of the constructs, which is determined solely by the chemical characteristics of the gel, in other words, by the cell-gel interaction. The sheets were extended to a 2 dimensional aggregate patterns in an attempt to create thick cellular sheets. Figure 12 shows two characteristic configurations, grid-like and close packed hexagonal arrangements. As expected, aggregates in the latter geometry fuse considerably faster, the emerging sheet has a uniform thickness overall (Jakab et al., 2004).
Fusion in Three Dimensions: Cellular Tube Formation.
From the fusion experiment in Figure 10 and 11 how the motility driven cell movements create shapes on the global scale can be observed. To study cellular rearrangements in the course of self assembly we embedded aggregates stained with red and green fluorescent dyes were embedded in 1.0 mg/ml collagen. Figure 13, A-3, shows the outcome of the fusion process after 80 h. Confocal imaging on the fused construct revealed that the emerging ring-like pattern fuses due to the movement of cells across the boundary of two adjacent aggregates, and not only to the adhesion between the cells on the aggregate's surface (Jakab et al., 2004).
The reason of building a tube like structure is because the tubular is the basic structure of all organs design. A living organism contains a variety of tubular structures with different cellular compositions and functions, as vasculature, intestines, lung, kidney, etc. Based on the optimal results on toroid formation, the scientist built a short tube from CHO cells consisting of three layers of 10 aggregates each, embedded in 1.0 mg/ml collagen gel. Different fluorescent staining of the layers was used to visualize the fusion in the vertical direction. Figure 13B shows brightfield and fluorescent images of the evolution of the tubular construct. Beside the radial contraction of the tube in Figure10, time lapse imaging revealed the axial shrinking, which ceases after 48 h (Jakab et al., 2004).
Scaffold Tissue Engineering.
Introduction of Scaffolds Tissue Engineering.
Scaffold technique of tissue engineering is different from the bio-printing whereby in scaffolding technique, cells are often implanted or 'seeded' into an artificial structure capable of supporting three-dimensional tissue formation. These structures, typically called scaffolds, are often critical, both ex vivo as well as in vivo, to recapitulating the in vivo milieu and allowing cells to influence their own microenvironments (Wikipedia, 2009a).
To achieve the goal of tissue reconstruction, scaffolds must meet some specific requirements. A high porosity and an adequate pore size are necessary to facilitate cell seeding and diffusion throughout the whole structure of both cells and nutrients. Biodegradability is often an essential factor since scaffolds should preferably be absorbed by the surrounding tissues without the necessity of a surgical removal. The rate at which degradation occurs has to coincide as much as possible with the rate of tissue formation. This means that while cells are fabricating their own natural matrix structure around themselves, the scaffold is able to provide structural integrity within the body and eventually it will break down leaving the neotissue, newly formed tissue which will take over the mechanical load. Injectability is also important for clinical uses (Wikipedia, 2009a).
Therefore the usage of the scaffolds tissue engineering can be applied into many fields such as bone and cartilage regeneration. Whereby it is crucial to study the basic of the formation of these bone and cartilage cells in order to for the tissue engineering to take place. Where this tissue engineering will have the advance technology to regenerate or reconstruct our broken bone or any part of the organ (Wikipedia, 2009a).
Scaffolds in tissue engineering of bone and cartilage.
Bone and cartilage generation by autogenous cell/tissue transplantation is one of the most promising techniques in orthopedic surgery and biomedical engineering. Whereby, autotransplantation is the transplantation of organs, tissues or even proteins from one part of the body to another in the same individual. Tissue transplanted by such "autologous" procedure is referred to as an autograft or autotransplant. It is contrasted with xenotransplantation (from other species) and allotransplantation (from other individual of same species) (Wikipedia, 2009a).
Bone and cartilage generation by autogenous cell/tissue transplantation is one of the most promising techniques in orthopedic surgery and biomedical engineering. Treatment concepts based on those techniques would eliminate problems of donor site scarcity, immune rejection and pathogen transfer. Osteoblasts, chondrocytes and mesenchymal stem cells obtained from the patient's hard and soft tissues can be expanded in culture and seeded onto a scaffold that will slowly degrade and resorb as the tissue structures grow in vitro and/or in vivo (Langer and Vacanti, 1993).
The scaffold or three-dimensional (3-D) construct provides the necessary support for cells to proliferate and maintain their differentiated function, and its architecture defines the ultimate shape of the new bone and cartilage. Several scaffold materials have been investigated for tissue engineering bone and cartilage including hydroxyapatite (HA), poly (a-hydroxyesters), and natural polymers such as collagen and chitin. Several reviews have been published on the general properties and design features of biodegradable and bioresorbable polymers and scaffolds (Chaignaud et al., 1997).
Polymer-based scaffold materials.
The meaning and definition of the words biodegradable, bioerodable, bioresorbable and bioabsorbable (Table 1) which are often used misleadingly in the tissue engineering literature are of importance to discuss the rationale, function as well as chemical and physical properties of polymer-based scaffolds. In this review, the polymer properties are based on the definitions given by Vert (Vert et al., 1992).
The tissue engineering program for bone and cartilage in the author's multidisciplinary research curriculum has been classified into six phases (Table 2). Each tissue engineering phase must be understood in an integrated manner across the research program from the polymer material properties, to the scaffold micro- and macro- architecture, to the cell, to the tissue-engineered transplant, to the host tissue. The first stage of tissue engineering bone or cartilage begins with the design and fabrication of a porous 3-D scaffold, the main topic of this review paper. In general, the scaffold should be fabricated from a highly biocompatible material which does not have the potential to elicit an immunological or clinically detectable primary or secondary foreign body reaction (Hutmacher et al., 1998).
Furthermore, a polymer scaffold material has to be chosen that will degrade and resorb at a controlled rate at the same time as the specific tissue cells seeded into the 3-D construct attach, spread and increase in quantity (number of cells/per void volume) as well as in quality (Wong and Mooney, 1997).
Table 1: Definition given by Vert.
Solid polymeric materials and devices which break down due to macromolecular degradation with dispersion in vivo but no proof for the elimination from the body (this definition excludes environmental, fungi or bacterial degradation). Biodegradable polymeric systems or devices can be attacked by biological elements so that the integrity of the system, and in some cases but not necessarily, of the macromolecules themselves, is affected and gives fragments or other degradation by-products. Such fragments can move away from their site of action but not necessarily from the body.
Solid polymeric materials and devices which show bulk degradation and further resorb in vivo; i.e. polymers which are eliminated through natural pathways either because of simple filtration of degradation by-products or after their metabolization. Bioresorption is thus a concept which reflects total elimination of the initial foreign material and of bulk degradation by-products (low molecular weight compounds) with no residual side effects. The use of the word & 'bioresorbable' assumes that elimination is shown conclusively.
Solid polymeric materials or devices, which show surface degradation and further, resorb in vivo. Bioerosion is thus a concept, too, which reflects total elimination of the initial foreign material and of surface degradation by-products (low molecular weight compounds) with no residual side effects.
Solid polymeric materials or devices, which can dissolve in body fluids without any polymer chain cleavage or molecular mass decrease. For example, it is the case of slow dissolution of water-soluble implants in body fluids. A bioabsorbable polymer can be bioresorbable if the dispersed macromolecules are excreted.
Table 2: The research program for tissue engineering bone and cartilage classified into six phases.
- Fabrication of bioresorbable scaffold
- Seeding of the osteoblasts/chondrocytes populations into the polymeric scaffold in a static culture (petri dish)
- Growth of premature tissue in a dynamic environment (spinner flask)
- Growth of mature tissue in a physiologic environment (bioreactor)
- Surgical transplantation
- Tissue-engineered transplant assimilation/remodeling
The design and fabrication of scaffolds in tissue engineering research is driven by three material categories:
- Regulatory approved biodegradable and bioresorbable polymers (Table 3), such as collagen, polyglycolide (PGA), polylactides (PLLA, PDLA), polycaprolactone (PCL), etc.
- A number of non-approved polymers, such as polyorthoester (POE), polyanhydrides, etc. which are also under investigation.
- The synthesis of entrepreneurial polymeric biomaterials, such as poly (lactic acidco-lysine), etc., which can selectively shepherd specific cell phenotypes and guide the differentiation and proliferation into the targeted functional premature and/or mature tissue.
In general, polymers of the poly (a-hydroxy acids) group undergo bulk degradation. The molecular weight of the polymer commences to decrease on day one (PGA, PDLA) or after a few weeks (PLLA) upon placement in an aqueous media (Wong and Mooney, 1997).
However, the mass loss does not start until the molecular chains are reduced to a size which allows them to freely diffuse out of the polymer matrix. Whereby this phenomenon is described and analyzed in detail by a number of researchers, results in accelerated degradation and resorption kinetics until the physical integrity of polymer matrix is compromised. The mass loss is accompanied by a release gradient of acidic by-products (Kronenthal, 1975).
In vivo, massive release of acidic degradation and resorption by-products results in inflammatory reactions, as reported in the bioresorbable device literature. If the capacity of the surrounding tissue to eliminate the by-products is low, due to the poor vascularization or low metabolic activity, the chemical composition of the by-products may lead to local temporary disturbances. One example of this is the increase of osmotic pressure or pH manifested through local fluid accumulation or transient sinus formation from fiber reinforced polyglycolide pins applied in orthopedic surgery (Bostmann et al., 1990).
Potential problems of biocompatibility in tissue engineering bone and cartilage, by applying degradable, erodable, and resorbable polymer scaffolds, may also be related to biodegradability and bioresorbability. Therefore, it is important that the 3-D scaffold/cell construct is exposed at all times to sufficient quantities of neutral culture media, especially during the period where the mass loss of the polymer matrix occurs (Bostmann et al., 1990).
The incorporation of a tricalciumphosphate (TCP), hydroxyapatite (HA) and basic salts into a polymer matrix produces a hybrid/composite material. These inorganic fillers allow to tailor the desired degradation and resorption kinetics of the polymer matrix. A composite material would also improve biocompatibility and hard tissue integration in a way that ceramic particles, which are embedded into the polymer matrix, allow for increased initial flash spread of serum proteins compared to the more hydrophobic polymer surface. In addition, the basic resorption products of HA or TCP would buffer the acidic resorption by-products of the aliphatic polyester and may thereby help to avoid the formation of an unfavorable environment for the cells due to a decreased pH (Agrawal and Athanasiou, 1997).
Control of the hydrodynamic and biochemical environment is essential for the successful in vitro engineering of 3-D scaffold/tissue constructs for potential clinical use (Freed et al., 1994).
Computer-controlled bioreactors that continuously supply physiological nutrients and gases, serve to regulate the required cell/tissue culture conditions for a long period of time. After the in vitro culturing of the 3-D scaffold/tissue construct, the degree of remodeling and cell/tissue replacement of the bone/cartilage transplant by the host tissue has to been taken into consideration. Cell and tissue remodeling is important for achieving stable mechanical conditions and vascularization at the host site. Hence, the 3-D scaffold/tissue construct should maintain sufficient structural integrity during the in vitro and/or in vivo growth and remodeling process. The degree of remodeling depends on the host anatomy and physiology. The polymer selection from a material science point of view is based on two different strategies in regard to the overall function of the scaffold. (Young et al., 1997.)
In the first strategy (Figure 14), the physical scaffold structure supports the polymer/cell/tissue construct from the time of cell seeding up to the point where the hard tissue transplant is remodeled by the host tissue. In the case of load-bearing tissue such as articular cartilage and bone, the scaffold matrix must serve an additional function. It must provide sufficient temporary mechanical support to withstand in vivo stresses and loading. In Strategy I research programs, the material must be selected and/or designed with a degradation and resorption rate such that the strength of the scaffold is retained until the tissue engineered transplant is fully remodeled by the host tissue and can assume its structural role (Hillsley and Frangos, 1994).
Bone is able to remodel in vivo under so-called physiological loading. It is a requirement that the degradation and resorption kinetics have to be controlled in such a way that the bioresorbable scaffold retains its physical properties for at least 6 months which include of 4 months for cell culturing and 2 months in situ. Thereafter, it will start losing its mechanical properties and should be metabolized by the body without a foreign body reaction after 12 -18 months (Figure 14) (Hillsley and Frangos, 1994).
The mechanical properties of the bioresorbable 3-D scaffold/tissue construct at the time of implantation should match that of the host tissue as closely as possible. It should posses a sufficient strength and stiffness to function for a period until in vivo tissue ingrowth has replaced the slowly vanishing scaffold matrix (Hillsley and Frangos, 1994).
A studied had done, whereby a poly (D,L-lactide-coglycolide) matrix under cyclic compressive loading. They concluded that changes in surface deformation and morphology suggest that the compressive loading initially collapses and stiffens the polymer matrix. The decrease in molecular weight is slowed down due to the reduction of surface area from hydrolysis, until the matrix architecture no longer accommodates the mechanical loading and begins to lose its integrity (Thompson et al., 1996).
For Strategy II (Figure 15), the intrinsic mechanical properties of the scaffold architecture templates the cell proliferation and differentiation only up to the phases, where the premature bone or cartilage is placed in a bioreactor. The degradation and resorption kinetics of the scaffold are designed to allow the seeded cells to proliferate and secrete their own extracellular matrix in the static and dynamic cell seeding phase (weeks 1 - 12), while the polymer scaffold gradually vanishes leaving sufficient space for new cell and tissue growth. The physical support by the 3-D scaffold is maintained until the engineered tissue has sufficient mechanical integrity to support itself (Thompson, et al. 1996).
Different research groups have shown in a number of studies that a nonwoven mesh made of polyglycolide fibers offers degradation and resorption kinetics for Strategy II. However, the challenge for the grown cell/tissue construct is to have similar mechanical properties to the host bone and cartilage. The research done by Ma and Langer, (1999) showed that cartilage which was cultured for 7 month in a bioreactor reached 40% of the mechanical properties of natural cartilage (Ma and Langer, 1999).
In an in vivo model, one of the major problems from a biomechanical and clinical view point is the primary mechanical stabilization of cartilage transplants (Thompson, et al. 1996).
Skeletal tissue, such as bone and cartilage is usually organized into 3-D structures in the body. For the repair and generation of hard and ductile tissue, such as bone, scaffolds need to have a high elastic modulus in order to be retained in the space they were designated for and also provide the tissue with adequate space for growth (Brekke, 1996).
If the 3-D scaffold is used as a temporary load-bearing device (Strategy II), the mechanical properties would maintain that load for the required time without showing symptoms of fatigue or failure. Therefore, one of the basic problems from a scaffold design point of view is that to achieve significant strength the scaffold material must have sufficiently high interatomic and intermolecular bonding, but must have at the same time a physical and chemical structure which allows for hydrolytic attack and breakdown (Brekke, 1996).
For tissue engineering a bone transplant, the creation of a vascularized bed ensures the survival and function of seeded cells, which have accessed to the vascular system for nutrition, gas exchange, and elimination of by-products. The vascularization of a scaffold may be compromised by purely relying on capillary ingrowth into the interconnecting pore network from the host tissue. In situ, the distance between blood vessels and mesenchymal cells are not larger than 100 lm. Therefore, the time frame has to be taken into account for the capillary system to distribute through larger scaffold volume. It may also be possible to control the degree and rate of vascularization by incorporating angiogenic and anti-angiogenic factors in the degrading matrix of the scaffold (Reece and Patrick Jr, 1998).
From a biomechanical and clinical point of view, the tissue-engineered bone or cartilage transplant should allow for a mechanically secure and stable fixation on or to the host tissue. For bone, the currently available medical devices, such as pins, screws, and plates might be used. However, the integration of a device-like part into the 3-D scaffold design can be advantageous (Dunkelman et al., 1995).
A number of fabrication technologies have been applied to process biodegradable and bioresorbable materials into 3-D polymeric scaffolds of high porosity and surface area. The conventional techniques for scaffold fabrication include fiber bonding, solvent casting, particulate leaching, membrane lamination and melt molding (Table 4). Several papers have reviewed the past and current research on scaffold fabrication techniques (Widmer and Mikos, 1998).
However for these researchers, they did not directly compared the 3-D scaffold-processing technologies for the tissue engineering community. Therefore, the aim of this part is to aggregate the compiled information and to present this data in a comprehensive form. Table 4 summarizes the key characteristics and parameters of the techniques currently used. In overall the process of tissue engineering bone and cartilage is not only to design, but also to fabricate reproducible bioresorbable 3-D scaffolds, which are able to function for a certain period of time under load-bearing conditions (Mikos et al., 1994).
Solvent casting, in combination with particle leaching, works only for thin membranes or 3-D specimens with very thin wall sections. Otherwise, it is not possible to remove the soluble particles from within the polymer matrix (Mikos et al., 1994).
Then by using the technology as describe above Mikos and his co-worker, fabricated porous sheets and laminated them to 3-D structures. Chloroform was used on the attachment interface for the lamination process. This fabrication technology is time consuming because only thin membranes can be used (Mikos et al., 1993).
Another disadvantage is that the layering of porous sheets allows only a limited number of interconnected pore networks. Solvent-casted polymersalt composites have also been extruded into a tubular geometry. The disadvantages of the above technologies include the extensive use of highly toxic solvents, time required for solvent evaporation, the labor intensive fabrication process, the limitation to thin structures, residual particles in the polymer matrix, irregularly shaped pores, and insufficient interconnectivity (Widmer et al., 1998).
The supercritical fluid-gassing process has been known for many years in the non-medical polymer industry as well as in the pharmaceutical community. This technology is used to produce foams and other highly porous products. The polymers which can be used for this technology have to have a high amorphous fraction. The polymer granules are plasticized due to the employment of a gas, such as nitrogen or carbon dioxide, at high pressures. The diffusion and dissolution of the gas into the polymer matrix results in a reduction of the viscosity, which allows the processing of the amorphous bioresorbable polyesters in a temperature of 30 - 40°C. The supercritical fluid-gassing technology allows the incorporation of heat sensitive pharmaceuticals and biological agents. However, on average only 10-30% of the pores are interconnected (Michaeli and Seibt, 1995).
In this case, Harris and his co-workers came out with an idea where they combined this technology with particulate leaching to gain a highly interconnected void network. The researchers could control porosity and pore size by varying the particle/polymer ratio and particle size (Harris et al., 1998).
The method used by Harris et al. (1998) wasn't good enough, therefore Whang and his co-workers came out with a protocol for the fabrication of aliphatic polyester-based scaffolds by using the emulsion freeze-drying method. Scaffolds with porosity greater than 90%, median pore sizes ranging from 15 to 35 µm with larger pores greater than 200 µm were fabricated. The scaffold pore architecture was highly interconnected which is necessary for tissue ingrowth and regeneration (Whang et al., 1999).
Based on their results from an animal experiment, the interdisciplinary group proposed a scaffold design concept which results in in-vivo bone regeneration based on hematoma stabilization (Whang K. et al., 1999).
The authors compare their in vivo bone engineering concept to the induction phase of fracture healing. The osteoprogenitor cells which are in the blood of the osseous wound are embedded in the scaffold microarchitecture via the hematoma. The multipotent cells differentiate to osteoblasts due to the presence of growth factors which are released by the host bone. However, this freezing-drying method are too technique sensitive whereby it is hard to control by the user. The fabrication of a truly interconnecting pore structure depends on the processing method and parameters as well as on the used equipment. (Zhang and Ma, 1999)
A group of two researchers had studied thermally induced phase separation technology to process polymeric 3-D scaffolds. This technique has been used previously to fabricate synthetic membranes for non-medical applications. The method has been extensively applied in the field of drug delivery to fabricate microspheres, which allows the incorporation of pharmaceutical and biological agents, such as bone morphogenetic proteins (BMPs) into the polymer matrix. In general, the micro and macro-structure is controlled by varying the polymer material, polymer concentration, quenching temperature, and solvents. However, this method is similar to the freezing-drying technique, therefore it is user and technique sensitive other than that the processing parameter had to be well controlled (Nam and Park, 1999).
A number of textile technologies have the potential to design and fabricate highly porous scaffolds. But there are only so-called non-woven mesh-like designs have been used to tissue engineer bone and cartilage. And excellent results in tissue engineering cartilage have been achieved by using non-woven meshes composed of polymer fibers of PGA, PGA/PDLA, and PGA/PLLA. In general, non-woven constructs can be only used for Strategy II since their physical properties do not allow load-bearing applications (Freed and Vunjak-Novakovic, 1998).
All the above-described technologies except the membrane-lamination method, do not allow the fabrication of a 3-D scaffold with a varying multiple layer design. Therefore, rapid prototyping technologies as well as so-called 'wafer stacking systems' shown in Figure 16, have the potential to design a 3-D construct in a multi-layer design within the same gross architectural structure (Hutmacher et al., 2004).
In engineering literature, Rapid prototyping Technologies (RP) also called Solid Free Form fabrication (SFF) methods are defined as a set of manufacturing processes that are capable of producing complex-free form parts directly from a computer-aided design (CAD) model of an object without part specific tooling or knowledge. Unlike machining processes such as milling, drilling, which are subtractive in nature, RP systems join together liquid, powder and sheet materials to form parts. Layer by layer, RP machines fabricate plastic, wood, ceramic and metal objects using thin horizontal cross sections directly from a computer-generated model (Beaman, 1997).
Rapid prototyping technologies, such as 3-D printing (3-DP) and fused deposition modeling (FDM) (Figure 17) allow the development of manufacturing processes to create porous scaffolds that mimic the microstructure of living tissue. Three-dimensional printing developed at the Massachusetts Institute of Technology is also a rapid prototyping technology which has been used to process bioresorbable scaffolds for tissue engineering applications (Park et al., 1998).
The technology is based on the printing of a binder through a print head nozzle onto a powder bed, with no tooling required. The part is built sequentially in layers. The binder is delivered to the powder bed producing the first layer, the bed is then lowered to a fixed distance, powder is deposited and spread evenly across the bed, and a second layer is built. This is repeated until the entire part, e.g. a porous scaffold, is fabricated. Following treatment, the object is retrieved from the powder bed and excess unbound powder is removed. The speed, flow rate and even drop position can be computer controlled to produce complex 3-D objects (Sittinger et al., 1996).
This printing technique permits CAD and custom-made fabrication of bioresorbable hybrid scaffold systems. The entire process is performed under room-temperature conditions. Hence, this technology has great potential in tissue engineering applications. Biological agents, such as cells, growth factors, etc., can be incorporated into a porous scaffold without inactivation if non-toxic binders, e.g. water can be used (Wu et al., 1996).
The FDM process forms 3-D objects from a CAD file as well as digital data produced by an imaging source such as computer tomography (CT) or magnetic resonance imaging (MRI). The process b