BioMEMS technology was originally based on the same technologies as MEMS. These technologies enabled low-cost, high-functionality devices in computer chip areas [1, 2]. The micro-fabrication of silicon-based structures usually includes processes such as photolithography, etching and deposition, to produce the desired configuration of features. In this way, BioMEMS structures that combine biologically sensitive elements with physical/chemical sensors can be developed to detect specific compounds in a given environment. Such BioMEMS sensors have been used to detect cells, proteins, DNA/RNA, or specific ions [3-5]. The digital capabilities of BioMEMS also allow them to precisely control and monitor local conditions for diagnostics applications. In general, the advantages of micro- and nano-scale detection technologies can be summarized as: (a) reducing the sensor elements to the scale of the target species, e.g. the size of a single cell, and hence take advantage of enhanced sensitivity and reduced detection time; (b) reduced reagent volumes and associated costs; and (c) integrating sample handling, mixing, separation, and detection for portability and miniaturization of the entire system. As for the development of BioMEMS devices, polymer-based devices and soft lithography have been increasingly attractive for use in biomedical applications. This is due to their increased biocompatibility, low cost, ability to integrate functional hydrogels, and rapid fabricating techniques [6-10]. Rapid response micro-fluidic channels  and pH-responsive micro-valves controlled by multiphase laminar flow  have also been successfully demonstrated. Just like MEMS are now considered as the technologies to interface the macro world to the micro/nano world, BioMEMS can play an important role for rapid and real-time analysis of cellular components, especially for single cells. Micro- and nano-scale systems and sensors enable precise measurements of protein, DNA, and chemical profiles for living cells in real time with controlled stimulus. These are essential for increasing our understanding of the underlying causes of basic cell functions such as differentiation, reproduction, apoptosis, etc., and their implications on various disease states. Spinal cord neurons for toxicological evaluation have been demonstrated on a microelectrode array via monitoring the microelectrode recordings [13, 14]. Gray et al.  fabricated cell based biosensors for toxicity monitoring, which resolved small extracellular potentials from embryonic chick cardiac myocytes. Powers et al.  studied 3D tissue structures of primary rat hepatocytes organized for 2 weeks in a microfabricated bioreactor with array of channels and cell-adhesive walls. Highly sensitive bio-chips have also been fabricated for measuring single ion channels within cell membranes , as well as specific proteins within living cells . These integrate sensors for the detection of DNA/RNA, proteins and other parameters. Such integrated systems make BioMEMS structures powerful platforms for micro/nano-scale cell studies.
Drug Delivery Methods and Thermo-responsive Polymers
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With the development of nanotechnologies, nanoparticles are promising candidates for the localized delivery of drugs to cancer cells , or the delivery of undeliverable molecules, such as RNA interfering effectors [20, 21]. The nanoparticles can also encapsulate macromolecular drugs to improve their bioavailability in the stomach, following oral administration. Furthermore, pH-responsive nanoparticles have been successfully applied for oral delivery systems based on the sharp pH gradient changes along the digestive tract [22-25]. Sajeesh et al.  fabricated poly(methacrylic acid) -chitosan-poly(ethylene glycol) nanoparticles that were loaded with insulin. The in vitro and in vivo experiments showed that nanoparticles displayed the expected insulin encapsulation efficiency and the release profile was largely dependent on the pH of the medium. Hu et al  fabricated chitosan-tripolyphosphate nano-particles for delivering tea catechins, and studies controlled release profile under different solution pH value. Metal nanoparticles [19, 27], which can be heated with external energy, have also been fabricated and applied for cancer targeting and local heating treatment. Transdermal delivery is an effective and alternative way when drugs will loose their effect during digestive processes, such as protein drugs. Among external drug delivery systems, microneedles are the most extensively developed technology for transdermal drug delivery [28-30]. The early microneedle systems were fabricated as solid needle arrays, which aimed at penetrating into skin to increase skin permeability for subsequent drug diffusion. In vivo experiments  demonstrated that such solid micro-needle can increase the permeability of human skin in vitro to calcein by up to 4 orders of magnitude. In order to improve the delivery capabilities of drugs in solution, hollow micro-needles were further developed based on solid micro-needles. The idea of perfusing drugs into tissues is similar to conventional injection. The hollow micro-needles were first fabricated in 2D arrays, i.e. all the needles lie on the support plane . However, the density of needles associated with the in-plane method is limited, since only one row of needles can be made per chip. Micro-needles perpendicular to the support plane were also fabricated . This configuration greatly enhances the delivery area. Micro-needles have been successfully used to deliver a variety of drug compounds into the skin, such as gene delivery  and vaccine delivery against Japanese encephalitis . Implantable micro-fabricated drug delivery systems enable localization of drug release and exposure at the target site, which totally avoid the poor bioavailability of protein drugs when orally administered. One widely employed design for internally implanted devices is micro-reservoir systems, which usually include reservoir array in a single device and can precisely control the openness of each single reservoir for drug release. The micro/nano fabrication technologies in MEMS provide powerful means to fabricate compact array of nano-liter reservoirs, usually in silicon based device. Santini et al  introduced a silicon microchip. This microchip can provide controlled release of single or multiple chemical substances on demand. Each dosage of drug is contained in a micro-reservoir that is covered with a gold membrane. The release mechanism is based on the electrochemical dissolution of the thin anode membranes. When applying anodic voltage to the membrane, electrochemical dissolution of the membrane is triggered in the presence of chloride ions. This leads to controlled membrane rupture, and then exposes the drug within the micro-reservoir to the surrounding tissue. Hence, the electrochemical dissolution of the anode gold membranes is the key process that enables the drug release from the micro-reservoir device. By combining the basic understanding of the electrochemical mechanism with electric circuit control, reliable and well-controlled device performance can be achieved. Pulsatile release of chemical substances with this device was demonstrated , i.e. small pulses of drug can be used to produce a complex release profile of multiple substances in order to maximize the effectiveness of drug therapies in the vicinity of the implants. For micro-reservoir chips, the standard drug loading methods include microinjection, inkjet printing, and micropipetting, etc. A rapid microwell filling method , which uses the principle of discontinuous dewetting, has also been reported. Two approaches can be used under this principle. One is to directly immerse the array in a bulk solution and then remove from the solution to fill the reservoirs. The other is to spread solution over the array to fill the wells and then allow extra liquid to drain off under gravity. This method allows uniform filling of the wells with small volumes down to ~3fL/well. Nanoporous silicon technology has also been applied for drug delivery purpose [39-41]. Silicon-based membranes consisting of arrays of uniform pores/channels show specific diffusion kinetics as the pore size down to nano-scale . By controlling the geometry of the nano-porous membrane, silicon nano-porous membranes can regulate the delivery kinetics of a wide range of drugs. In vivo experiments demonstrated that the diffusion from the nano-porous membrane effectively prolonged levels of bovine serum albumin in the blood . Moreover, for drug delivery systems utilizing nano-porous membranes, the release mechanism is attributable only to the constrained diffusion process, i.e. no moving parts (like mechanical pumps) are required. Such a nonmechanical driving mechanism offers important advantages in drug delivery applications that require high drug load ratios and flexibility with respect to the encapsulated drug state.
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As another drug delivery approach, thermo-responsive hydrogels that can change their swelling behavior and fluid release characteristics in response to environmental temperature have raised increasing interest in recent years [42-46]. The capacity to store and release drugs has made those hydrogels attractive candidates for applications in drug eluting systems [46, 47]. Within this context, extensive studies have been done in smart thermo-sensitive hydrogels that can release controlled amounts of drugs under well-defined temperature conditions [46-48]. Poly(N-isopropylacrylamide) (PNIPA) hydrogel, prepared from the monomer N-isopropylacrylamide (NIPAAm) and a suitable cross-linker, is a well-known thermo-responsive hydrogel that undergoes a lower critical solution temperature (LCST) around 32oC in aqueous solution . The 32oC LCST is close to physiological body temperature (37oC). This makes PNIPA very appealing for applications in controllable drug release [50-53]. As PNIPA hydrogels swell much larger at temperature lower than LCST, the extent and rate of the swelling can be controlled by temperature, without altering the chemical environment. Therefore, drug release profiles can be remarkably controlled by alternation of the solution temperature. However, the deswelling rate of conventional PNIPAAm hydrogel was observed to be very slow, owing to the formation of a dense and thick skin layer, which prevents water molecules from migrating out of the gel when the deswelling occurs [54, 55]. Hence, interpenetrating polymer networks (IPNs) have been introduced to form water-releasing channels within the PNIPAAm hydrogel network. This improves their swelling/deswelling performance [54-56]. Besides specific polymer matrix designs for hydrogel drug delivery, mathematical models of the drug release mechanism have also been developed for specific applications [57-59]. Fick's law is the most fundamental mechanism for describing drug release from hydrogels. Fick¡¯s first and second laws can describe the diffusion that occurs under constant (or time dependent) diffusion coefficients during diffusion-controlled drug release [42, 60]. However, it is usually hard to obtain analytical solutions to Fick's law when drug release with complex geometries or further modeling on diffusion coefficients is needed to describe complicated release mechanisms. For hydrogel drug delivery, the drug diffusion usually accompanies with large extent polymer matrix swelling. The swelling process, which usually happens together with phase change, can significantly affect the overall hydrogel release performance and lead to swelling-controlled release other than traditional Fickian diffusion. As an empirical equation developed by Peppas et al. [61, 62], a time-dependent power law function can describe the early stage of swelling diffusion processes. However, more sophistical modeling is required to solve the moving boundary conditions for the modeling of swellable hydrogels
Interfacial Adhesion Strength between Cell and Substrate
For implantable BioMEMS, the interfacial adhesion between biological cellular layers and substrates/devices is very important. Hence, there is a need for development of a fundamental understanding of cell-surface adhesion, and methods to enhance adhesion between biological tissues and the surface need to be developed. Due to the significance of the cell-surface adhesion, a number of researchers [32, 66-73] have developed different techniques for the measurement of adhesion (Figure2. 1). Atomic force microscopy (AFM) has been used to measure the ligand-receptor bonds associated with cell/surface adhesion (Figure 2.1a). By coating special proteins on AFM tip, adhesion forces with pico-Newton magnitudes have been measured for specific ligand-receptor interactins [74-76]. Zhang and Moy  have quantified the detachment force of streptavidin¨Cbiotin by the downward deflection of the cantilever. Li et al.  have also measured the dynamic response of ¦Á5¦Â1 integrin-fibronectin to a pulling force, and the single molecule rupture force was measured between live K562 cells and fibronectin. Huang et al.  have used a glass beam (75 ¦Ìm in diameter) as a cantilever to contact and detach single chondrocytes from glass slides, The detachment force after 6 hours of cell culture was reported to be 388 ¡À 78 nN. Another widely used method is micropipette aspiration technique [32, 66, 68]. This technique has been used widely to study the time-dependent deformation of living individual cells that are subjected to extracellular pressure, as shown in Figure 2.1b. The applied aspiration pressure ranges from 0.1-1000 Pa, with a resolution of 0.1 Pa [77, 78]. During the aspiration measurement, a single living cell is drawn into a glass tube, via stepwise application of aspiration pressure (i.e., suction). The inner diameter of the tube is a chosen fraction of the nominal diameter of the cell, and the aspiration pressure is maintained over a specified duration. The attendant extension of the cell into the pipette can be monitored via optical microscopy. The displacement of the cell membrane that is tracked by light microscopy was claimed with a resolution of 25 nm . Micropipette aspiration enables real-time correlation of pressure and whole-cell deformation, while several uncertainties arise in the mechanical property measurements. It is also difficult to keep cell in a proper position inside the micropipette, while the sharp tube edge causes artificial stress singularities during the aspiration of cells into micropipettes. The strong adhesion of cell membrane to the inner wall of the micropipette can also greatly affect the elastic properties measurements. Compared to micropipette aspiration experiments, the optical tweezers stretch method [71, 72] can provide much better force-displacement characteristics of single cells. In this technique, laser traps control the position of two silica beads, which are attached to opposite ends of a single cell, first to moving the laser traps, and therefore the silica beads, leading to the stretching of the cell (Figure 2.1c). This well-controlled direct tensile stretching of biological cells makes the direct interpretation of the experimental observations simpler. This method further makes it easier to investigate the progression of a disease status, such as those associated with the infestation by red blood cells with malaria parasite. Plasmodium falciparum has been shown to changes the elastic and viscoelastic properties of red blood cells [79, 80]. Supported by the development of micro-fabrication techniques, large area arrays of fabricated elastomeric pillars have been used successfully to measure the mechanical interactions between cells and their underlying substrates [81-83]. The deflections due to cell attachment and spreading can be controlled by the geometry of micro-posts. For small deflections, the deflection is directly proportional to the force applied by the attached cell. The deflection behavior of the posts is also well described by linearly elastic beam theory. Furthermore, since each pillar is independently deflected, the measured traction forces under the cells are direct and localized. The cell traction forces detected by miro-pillar-arrays are around 50 nN . The ability to directly detect sub-cellular distributions of traction forces makes micro-fabricated arrays attractive for cell mechanics studies. The soft hydrodynamic force exerted by fluid flow has also been used to study the cell-surface interactions on biomedical surfaces [67, 84, 85]. Under fluid flow conditions, the applied hydrodynamic forces are less likely to induce additional biological responses that may compromise the measured strength data. Hence, there has been increasing interest in the use of the shear assay method for the ¡°soft¡± detachment of cells from biomedical substrates. In early stage, centrifugation [67, 86] and rotating plates [87, 88] have been used widely to measure the adhesion strength. However, in-situ observations of living cells are difficult during such measurements. The parallel flow chamber method [84, 85, 89-92] has been further developed for the in-situ observation of cell deformation and detachment during (cell culture) fluid flow across cells in a micro-fluidic channel (Figure 2.2). Detailed cell adhesion strengths have been calculated using fluid flow conditions [84, 91, 93]. These include computational fluid dynamics simulations, e.g. Goldstein and DiMilla  simulated the detachment of murine 3T3 fibroblasts from self-assembled monolayers, and the simulation of the adhesion of circulating leukocytes [95-99] to the vascular endothelium. Research also revealed that shear stress can further change the subcellular structure and focal adhesion details [100-102].
Effects of Surface Micro-texture on Cell Alignment
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Surface chemistry has been shown to effectively improve cell adhesion [105-107]. The integrin and other receptors located within the cell membrane can interact with surface proteins, and a cascade of intracellular signaling events is triggered by the coated protein layer. Hence, protein coating on biomaterials surfaces enables the control of direct cell responses to foreign bodies. Several adhesion-related protein coatings have been examined. The most widely investigated of these protein coatings are the cell adhesive acid RGD , fibronectin , laminin , and collagen . Surface micro-texture has also been proven to be important and effective in the moderation of protein adsorption, cell-surface interactions, and host response to biomaterials. In recent years, significant efforts have been made to develop micro-textured polymeric materials for applications in biomedical systems [111-116]. In most cases, the micro-textures have been introduced to enhance tissue integration and wound healing [103, 114, 117-119] Numerous techniques have also been used to produce micro-textured surfaces. These range from micro-machining and grit blasting to more controlled fabrication methods, such as laser-texturing [120-122] and photolithography [117, 123-125]. In some cases, improved cell adhesion, greater physical integration of the device can be achieved via micro-texturing of surfaces in vivo environments. Among the various surface topologies, micro-grooves have been widely used. The use of micro-grooved geometries has been shown to promote ¡°contact guidance¡± [117, 126], which is a phenomenon that involves the alignment of cells, as they spread across micro-groove geometries. Such contact guidance has been shown to reduce capsule formation during tissue integration [103, 127]. Significant efforts of micro-grooves have been put on bioactive materials, such titanium  and Ti-6Al-4V [120, 128]. Laser-ablated micro-grooves have also been used to modify the surface characteristics of biomaterials and influence cellular behavior, such as contact guidance. Careful work [122, 128] has been done on the cell-surface interaction between human osteo-sarcoma (HOS) cells and laser-grooved Ti-6Al-4V. The laser micro-grooved geometries with controlled spacings and depths showed strong effects on cellular alignment of HOS cells in 2 day culture experiments. Immuno-fluorescence staining also revealed subcellular cytoskeleton (actin) and focal adhesion (vinculin) rearrangements. Poly-di-methyl-siloxane (PDMS) has also been widely used in microfluidic devices, BioMEMS and lab-on-a-chip applications [115, 129-131]. The increasing interest in PDMS has been due largely to its excellent combination of biocompatibility and mechanical properties [132-137]. Some researchers have also studied the effects of micro-groove geometry on cell spreading and alignment on PDMS [108, 109, 138-141]. The geometry of micro-groove, like groove height and spacing, has been proved as important factors that affect the cell alignment. Clark et al.  studied the effects of single crossing step on cell behaviors, using baby hamster kidney cells and Madin Darby canine kidney cells on grooved Perspex with varying dimensions . They showed that repeat spacing had a small effect when ranging in 4-24 ?m, but that groove depth (0.2-1.9 ?m) proved to be much more important in determining cell alignment. Compared to the Ti-6Al-4V with polished (smooth) and Al2O3 blasted (rough), enhanced orientation and attachment were observed on the Ti-6Al-4V micro-grooved surfaces with groove spacings of 20 ?m, depths of 10 ?m, and widths of 11 ?m . Walboomers et al.  have also studied the spreading of rat dermal fibroblasts on micro-grooved polystyrene substrates. Furthermore, with the advent of nanofabrication techniques, nano-scale groove were introduced in cell alignment studies. A number of experiments [144-147] have shown that nano-scale grooves have strong alignment effects on cell alignment, as well as focal adhesions. Cells were observed to elongate and align along nano-scale grooves and ridges, and width of focal adhesions is controlled by the ridge width. In general, micro- and nano-fabrication techniques offer effective approaches to the understanding of cell-surface interactions. They can also be used to directly modify cell behavior on implanted devices. In conclusion, the surface texturing of biomedical systems can be used to enhance adhesion and integration, reduce scar tissue formation, and moderate immune responses.