Microsystem Technologies

, Volume 12, Issue 1, pp 120–127

Bio-MEMS fabricated artificial capillaries for tissue engineering

Authors

    • Department of Mechanical EngineeringNational Chung-Hsing University
  • C. L. Chen
    • Department of Mechanical EngineeringNational Chung-Hsing University
  • S. H. Hsu
    • Department of Chemical EngineeringNational Chung-Hsing University
  • Y. L. Chiang
    • Department of Chemical EngineeringNational Chung-Hsing University
Technical paper

DOI: 10.1007/s00542-005-0017-7

Cite this article as:
Wang, G.J., Chen, C.L., Hsu, S.H. et al. Microsyst Technol (2005) 12: 120. doi:10.1007/s00542-005-0017-7

Abstract

In this report, we focus on the microfabrication and cell seeding issues of artificial blood capillaries for tissue engineering. Two different fabrication methods (stainless steel electroforming and silicon electroforming) and a number of materials (PC, Polycarbonate and biocompatible material PLGA, poly lactide-co-glycolides) are implemented to build the vascular network. The vascular network is then used as the scaffold to cultivate the bovine endothelial cell (BEC). During the period of cell cultivation, oxygen and nutrient need to be continuously delivered by a circular pressurizing system. In cell culture, encouraging results are obtained through the dynamical seeding of the BEC on the scaffolds. A systematic cell culture process has been developed after repeated experiments. Successful seeding efficiencies are obtained by using the developed systematic cell culture process.

1 Introduction

Most of the vital organs are short of the ability of regeneration. Organ transplant is the only solution to the loss or failure of a vital organ. Shortage of donor organs makes tissue engineering a promising solution to organ transplantation.

Tissue engineering is a new field in science, medicine, and engineering in which bio-artificial organs and tissues are built and then implanted into a live body to repair or increase the functions of the vital organs. The concept of tissue engineering is to cultivate cells in a biodegradable scaffold in which the cells grow to regenerate new organs such as skin, cornea, born, and cartilage. One of the major challenges of the tissue engineering is the lack of intrinsic blood vessels to transport the nutrient and metabolite. Once the tissue is larger then 1–2 mm, the cultivating cells will shrivel due to the lack of metabolic nutrient and air. It is thus desired to provide the artificial tissues with artificial blood vessels or capillaries.

In general, growth accelerant for the cells is spread on the biodegradable scaffold such that the cells can be catalyzed to regenerate blood vessels. However, blood capillaries with diameter in μm are difficult to be constructed by conventional approaches. King et al. (2002) and Borenstein et al. (2002) were the pioneers and primarily microfabricated capillary scaffold on the PolyDiMethylSiloxane (PDMS) and seeded with endothelial cells to form artificial capillaries. Since the PDMS is transparent, the cell cultivating processes can be observed. However, the scaffolds of the vessel network have to be made of a biodegradable material such that the endothelial cells can be successfully seeding and implanted into the body of animals.

In the past few years, bioMEMS technology such as bulk micromachining (Kovacs et al. 1998; Pan et al. 2002), LIGA processing (Chang and Kim 2000; Marques et al. 1997) hot embossing (Jeon and Chiu 2002; Becker and Heim 1999), and excimer laser cutting (Hui and Qin 2002; Kancharla and Chen 2002) have been successfully used to build microfluidic structures and devices such as the microvalve, the micropump, the microreactor, and the biochip in the range of submicron. Many useful biomedical devices have been developed, e.g. biochip and lab-on-a-chip (Madou et al. 2001; Martin et al. 2000; Jiang et al. 2000), biosensor (Polla et al. 2000; Alvarez et al. 2003), electrotactile display (Tang and Beebe 2003). It seems feasible to apply the BioMEMS technology to the manufacturing of the capillary network.

In this article, we focus on the development of the microfabrication method for biocompatible capillaries which are as small as the real capillaries and capable of being observed the cell seeding processes. The microfabrication process of the biocompatible PLGA scaffolds is the core technique of this new developing method. The procedures to micromachine the PLGA scaffolds are: (1) lithograph the silicon wafer or stainless steel substrate (2) process the electroforming to make the mold (3) spin coat the liquid PLGA that is dissolved by certain solvent on the mold (4) use O2 plasma to join the PLGA and the glass together.

2 Design and Fabrication of the Vascular Network

2.1 Vascular Network Design

Since the dimension of the endothelial cell is around 10–15 μm, capillary dimension on the vascular network is designed to be 20×60 μm2 with 800 μm in segment length. The included angle at each node is chosen to be 30° to avoid large pressure drop between vascular segments (Chang et al. 2000). Larger pressure drop may induce fluid circulation problem. A lower injecting pressure not only can prevent the possible crack of the microchannel from happening but also ensures the cells can stably adhere to the surface of the scaffold. Figure 1 illustrates the proposed vascular network.
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Fig. 1

Vascular network pattern

2.2 Vascular Network Fabrication

The fabrication processes consist of mode fabrication, replica molding, and bonding three steps. For comparisons, two different mode fabrication methods and materials are implemented to build the vascular network.

2.2.1 Replica Mode Fabrication

(1) Stainless steel electroforming method

Stainless steel electroforming processes are schematically illustrated in Fig. 2 and itemized as follows: (1) prepare the stainless steel substrate that has suitable surface roughness (2) spin coat thick-film photoresist SU-8 on the substrate (3) photo transfer the vascular network pattern (4) make Ni-Co electroforming (5) remove photoresist.
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Fig. 2

Stainless steel electroforming

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Fig. 3

Hydrogen bubble induced defects

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Fig. 4

Effect of the moisture preserver

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Fig. 5

Stainless steel based replica mode

During electroforming, the nickel and hydrogen reductions take place at the cathode simultaneously such as shown in Eq. 1. Structure defects such as shown in Fig. 3 due to hydrogen bubbles can be observed. If suitable moisture preserver is required so that the bubble induced defects can be eliminated (Fig. 4). After removing the photoresist and cleaning by DI-water, the replica mode is shown in Fig. 5.
$$\begin{aligned} \hbox{Ni}^{2 +} + 2e^ - \to \hbox{Ni}\& \\ 2\hbox{H}^ + + 2e^ - \to \hbox{H}_2\& \\ \end{aligned} $$
(1)

Stainless steel substrate possesses advantages such as high strength, high corrosion resistance, and can be directly used as conducting layer. Once the thickness of the photoresist is higher than that of the electroforming layer, the metal depositing rate can be controlled through the current density such that the desired electroforming thickness can be precisely met. The drawback is that the adhesion between the substrate and the electroforming layer is not strong enough when the surface of the substrate has not been well cleaned.

(2) Silicon electroforming method

Schematic illustration of the silicon electroforming processes is shown in Fig. 6. The detail procedures are: (1) spin coat the thick-film photoresist JSR on a clean silicon wafer (2) photo transfer the vascular network pattern (3) sputter a 0.1–0.2 μm thick silver conducting layer on the substrate (4) electroform Ni-Co alloy as the mode body (5) wet-etch the silicon wafer and remove photoresistor.
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Fig. 6

Silicon electroforming

The main advantage of the silicon electroforming method is that the semiconductor fabrication techniques can be easily adopted. However, residual stress that results from the relatively larger volume of the electroforming layer can be a serious problem. Nickel that has less hardness can be used to reduce the residual stress.

2.2.2 Hot Embossing and Bonding

In this research, the high penetrability PC (Polycarbonate) and the biodegradable PLGA (poly lactide-co-glycolides) polymer materials are selected as the substrates. The PC based vascular network is built to enable the cultivating processes to be observed.

(1) PC substrate

PC is a common material used in microfluidic applications. Both the formation from electroforming mold and the bonding of PC layers to complete the vascular network processes are conducted by using a hot embossing machine.
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Fig. 7

Embossing without preheating

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Fig. 8

Embossing without enough time

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Fig. 9

Vascular network after embossing

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Fig. 10

Surface profile of the embossed vascular network

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Fig. 11

Insufficient temperature and pressure bonding

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Fig. 12

Over-heating bonding

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Fig. 13

Adequate bonding

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Fig. 14

Apparatus for cell cultures

During the hot embossing, there are four control variables: preheating, hot embossing temperature, loading pressure, and embossing period are chosen to investigate their influences on the embossing results. It is found from experiments that preheating is crucial to the embossing quality. Figure 7 indicates the embossing result without preheating. The possible reason is that the PC substrate deformed under external loading at room temperature. It is also observed that a shorter embossing time (less than 3 min) results in unwanted microfluidic channels (Fig. 8). However, embossing temperature and loading pressure are the two key factors to the embossing quality.

In our experience, the following parameter set can be implemented to produce better vascular network patterns such as shown in Fig. 9; preheating time = 5 min, embossing temperature = 135°C, loading pressure = 4.5 Mpa, embossing time = 6–10 min. Figure 10 illustrates the surface profile of the vascular network measured by surface roughness metrology.

For the bonding of PC layers, bonding pressure and temperature are the main control variables. Figure 11 shows the bonding results of insufficient temperature and pressure. Figure 12 illustrates the effect of over-heating.

In our experiments, serviceable bonding as shown in Fig. 13 can be carried out by adequate parameter settings such as: preheating time = 5 min, bonding temperature = 137°C, bonding pressure = 5.6 Mpa, bonding time = 6–8 min.

(2) PLGA Substrate

PLGA is one of the biodegradable biomedical materials. It is a good candidate for making scaffolds for vascular network (King et al. 2002). The PLGA vascular network fabrication procedures are itemized as,
  1. 1

    PLGA substrate preparation: Dissolve the PLGA 50/50 (Birmingham Polymers, U.S., Mw ≈ 80,000) into Dioxane (TEDIA); Heat the solution up to 60°C and stir simultaneously to become a 20% PLGA solution.

     
  2. 2

    PVA spin coating (for PLGA demolding): Make the PVA solution with 1% concentration; Spin coat the PVA film on the replica mold.

     
  3. 3

    PLGA spin coating: Spin coat the PLGA film to cover the PVA film; Exhaust at room temperature for 24 hours to evaporate the solvent.

     
  4. 4

    Demolding: Dip the replica mold into DI water for a few hours to hydrolyze the PVA film; Demold to obtain the PLGA based scaffolds.

     

During cultivation, the PLGA scaffolds become opaque after being immersed in the culture solution for long-term culture. In this research, the PLGA substrate is bonded with glass plate by O2 plasma such that the cultivating processes can be observed and recorded.

3 Cell cultivation

In this work, the bovine endothelial cells (BEC) are adopted to seed on the vascular scaffolds. However, dynamic seeding with circulating flow is much more complicated than simple static seeding on cultivation containers. To gain a deeper understanding about the dynamic seeding on the vascular network, preliminary experiments using a straight microchannel with 20×150 μm2 are conducted. The experimental apparatus is schematically illustrated in Fig. 14.

3.1 Biocompatibility experiments

Biocompatibility is the key factor that determines the adhesion between cells and the scaffolds. Different kinds of microchannels, PC substrate, PC with CBD-RGD, and glass substrate are used for biocompatibility comparisons. The CBD-RGD process is to modify the cellulose binding domain (CBD) gene of the trichoderma koningii by polymerase chain reaction to increase the adhesion. Due to its high hydrophilic identity, the glass scaffold is used for contrast. Biocompatibility experimental results are shown in Fig. 15. After 12 h of seeding, the glass scaffold has the largest number of living cells, followed by the PC and PC+RGD. However, the PC+RGD scaffold overtakes the simple PC scaffold after 48 h of seeding. It is demonstrated that the CBD-RGD process increases the biocompatibility of the PC scaffold.
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Fig. 15

Biocompatibility experiments a 12 h of seeding b 48 h of seeding

3.2 Static Seeding

Detail static seeding procedures are listed below.
  1. 1.

    Sterilization: Firstly sterilize the scaffolds and devices by deep UV-light. Followed by a second sterilization by 70% ethanol that is injected through a peristaltic pump. The residual ethanol is then replaced by the buffer solution PBS.

     
  2. 2.

    CBD-RGD: Inject the RGD solution into the microchannel; Seal both outlets for 30 min to allow enough reacting time.

     
  3. 3.

    Cell suspension preparation: Prepare cell suspension with concentration 106–107 cells/ml.

     
  4. 4.

    Static seeding: Pump cell suspension into microchannel; Static seed inside the cultivating container under the conditions of 37°C, 5 CO2, relative humidity 95%.

     
In general, cells should stick on the surface of the scaffolds after 12–24 h of static seeding. When cells have stuck on the surface, dynamic seeding can be conducted. Figure 16 illustrates the static seeding processes for the 4×106 cells/ml cell suspension. After 6 h of static seeding, most of the cells have stuck on the surface. Six more hours later, cells gradually stretch out. However, the generality of cells are dead after 24 h of static seeding. It is presumed that the culture medium inside the microchannel is insufficient. Therefore, a semi-dynamic seeding trail with medium exchanged every 6 h is implemented. After semi-dynamic seeding for 3 days, almost all cells can adhere to the microchannel. In our later experiments, two days of semi-dynamic seeding is enough to ensure the cell adhesions. Dynamic seeding can thus be conducted.
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Fig. 16

Static seeding for the 4×106 cells/ml cell suspension a Initial stage b Six hours in seeding c Twelve hours in seeding

3.3 Dynamic Seeding

The apparatus for dynamic seeding is shown in Fig. 14, in which the culture medium is continuously circulated by the peristaltic pump so that the environment of a real artery can be imitated. Since the peristaltic pump is set outside the culture container, little crack on the pipe junction may induce serious infection to the seeding cells. It should be carefully avoided.

Figure 17 shows the seeding processes of the cell suspension with a concentration of 107 cells/ml. After 12 h of static seeding, most of the cells stick on the surface. Semi-dynamic seeding carries on for 48 h followed by dynamic seeding. After 96 h of dynamic seeding at a circulating rate of 8 μl/min, it is observed that cells completely adhere to the microchannel toward near-confluence. It is also found that the pressure variations induced by the bubbles in the micro-channels will cause serious damages on the already adhered cells; therefore, the bubble effects in the micro-channels need to be eliminated as far as possible.
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Fig. 17

Cells seeding for the 4×107 cells/ml cell suspension a Initial stage b 12 h in seeding (static) c 48 h in seeding (semi-dynamic) d 96 h in seeding (dynamic)

For the more complicated and smaller vascular network, a smaller circulating rate is required to prevent the microchannels from cracking. Figure 18 depicts the culture images of the 4×107 cells/ml cell suspension with a medium circulated at 2 μl/min. After 8 h in culture, cells begin their attachment. For 48 h in culture, cells attach completely and tightly stick together.
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Fig. 18

Illustrations of the vascular network cells seeding a Initial stage b 8 h in seeding c Cells after 48 h

4 Conclusions

The main goal of this research is to use the BioMEMS technique to build artificial blood capillaries on both the PC (Polycarbonate) and biocompatible material such as PLGA (poly(lactide-co-glycolides)). Firstly, a vascular network is constructed on both materials. The micro-channel network is then used as the scaffold to cultivate bovine endothelial cells (BECs). The artificial blood capillaries are finally constructed after the endothelial cells grow up and the scaffold is decomposed. During the period of cell cultivation, oxygen and nutrients need to be continuously administered by a circular pressurizing system.

In cell culture, encouraging results are obtained through the dynamical seeding of the BEC on the PC based scaffolds. A systematic cell culture process has been developed after repeated experiments. Successful seeding efficiencies are obtained by using the developed systematic cell culture process.

Acknowledgements

The authors would like to address their gratitude to the National Science Council of Taiwan for financial support under grant NSC-91-2212-E-005-012. The work was conducted in the Center of Tissue Engineering and Stem Cells Research (TESC) at the National Chung-Hsing University, Taiwan. The center is funded by National Health Research Institutes (NHRI).

Copyright information

© Springer-Verlag 2005