Microfabricated drug delivery systems: from particles to pores

Advanced Drug Delivery Reviews 55 (2003) 315–328 www.elsevier.com / locate / addr

Microfabricated drug delivery systems: from particles to pores Sarah L. Tao a,b , Tejal A. Desai a,b , * a

Department of Bioengineering, University of Illinois at Chicago, 851 S. Morgan Street, Chicago, IL 60607, USA b Department of Biomedical Engineering, Boston University, 44 Cummington Street, Boston, MA 02215, USA

Abstract Microfabrication techniques which permit the creation of therapeutic delivery systems that possess a combination of structural, mechanical, and perhaps electronic features may surmount challenges associated with conventional delivery of therapy. In this review, delivery concepts are presented which capitalize on the strengths of microfabrication. Possible applications include micromachined silicon membranes to create implantable biocapsules for the immunoisolation of pancreatic islet cells—as a possible treatment for diabetes—and sustained release of injectable drugs needed over long time periods. Asymmetrical, drug-loaded microfabricated particles with specific ligands linked to the surface are proposed for improving oral bioavailability of peptide (and perhaps protein) drugs. In addition, microfabricated drug delivery systems ranging from transdermal microneedles to implantable microchips will be discussed.  2002 Elsevier Science B.V. All rights reserved. Keywords: Silicon; Microtechnology; Microfabrication; Therapeutic; Drug delivery; Microparticles

Contents 1. 2. 3. 4.

Introduction ............................................................................................................................................................................ Controlled release drug delivery systems................................................................................................................................... Microfabrication technology .................................................................................................................................................... Microneedles for transdermal drug delivery............................................................................................................................... 4.1. Microfabrication of silicon microneedle arrays ................................................................................................................... 4.2. Transdermal transport studies............................................................................................................................................ 4.3. Other microneedles .......................................................................................................................................................... 5. Implanted microchip for localized drug delivery ........................................................................................................................ 5.1. Microchip design ............................................................................................................................................................. 5.2. Irreversible metallic valves ............................................................................................................................................... 5.3. Reversible polymeric valves ............................................................................................................................................. 5.4. Current developments....................................................................................................................................................... 6. Bioadhesive microparticles for oral drug delivery ...................................................................................................................... 6.1. Silicon dioxide microparticles ........................................................................................................................................... 6.1.1. Fabrication of microparticle body ............................................................................................................................ 6.1.2. Fabrication of microparticle reservoir ...................................................................................................................... 6.1.3. Surface modification chemistry ............................................................................................................................... 6.2. Poly(methyl methacrylate) microparticles........................................................................................................................... *Corresponding author. Tel.: 1 1-617-358-3054; fax: 1 1-617-353-6766. E-mail address: [email protected] (T.A. Desai). 0169-409X / 02 / $ – see front matter  2002 Elsevier Science B.V. All rights reserved. doi:10.1016/S0169-409X(02)00227-2

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6.2.1. Fabrication of microparticle body ............................................................................................................................ 6.2.2. Fabrication of microparticle reservoir ...................................................................................................................... 6.2.3. Surface modification chemistry ............................................................................................................................... 6.3. Release mechanism .......................................................................................................................................................... 6.4. Microparticle bioadhesion to intestinal epithelium .............................................................................................................. 6.4.1. Lectins .................................................................................................................................................................. 6.4.2. Caco-2 binding ...................................................................................................................................................... 6.5. Future developments ........................................................................................................................................................ 7. Nanoporous immunoisolating biocapsules ................................................................................................................................. 7.1. Allotransplantation ........................................................................................................................................................... 7.2. Biocapsule design ............................................................................................................................................................ 7.3. Biocompatibility and cytotoxicity studies ........................................................................................................................... 8. Conclusions ............................................................................................................................................................................ Acknowledgements ...................................................................................................................................................................... References ..................................................................................................................................................................................

1. Introduction The application of micro- and nanotechnology to the biomedical arena has tremendous potential in terms of developing new diagnostic and therapeutic modalities. Over the last several years, microfabrication technology has been applied to the successful development of a variety of health care-related products including diagnostic (‘lab-on-a-chip’) systems and techniques and apparatus for high throughput screening of new drug candidates [1,2]. While the majority of research has focused on the development of miniaturized diagnostic tools, researchers have more recently concentrated on the development of microdevices for therapeutic applications. Microand nanofabrication techniques are currently being used to develop implants that can record from, sense, stimulate, and deliver to biological systems. Micromachined neural prostheses, drug delivery micropumps, tissue scaffolds, and stents [3–6] have all been fabricated using precision-based microtechnologies. Drug delivery remains an important challenge in medicine [7] and microfabrication techniques may be used to develop novel drug delivery devices with capabilities not possible with current systems. This paper will review some of the current and future approaches that utilize microfabrication technology for drug delivery.

2. Controlled release drug delivery systems Conventional dosage forms, such as oral delivery and injection, are the predominant routes for drug

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administration. However, these types of dosages are not easily able to control the rate of drug delivery or the target area of the drug and are often associated with an immediate or rapid drug release. Consequently, the initial concentration of the drug in the body peaks above the level of toxicity and then gradually diminishes over time to an ineffective level. The duration of therapeutic efficacy then becomes dependent on the frequency of administration, and half-life of the drug, and high dosages of non-targeted drugs are often administered to achieve an effective blood concentration [8]. In recent years, increasingly sophisticated and potent drugs have been developed by the biotech industry. For many of these new protein-based and DNA-based compounds, the therapeutic concentration range is often small, toxicity is observed for concentration spikes, or the therapeutic concentration range varies with time, which renders traditional methods of drug delivery ineffective [9]. An immense amount of interest has been increasingly placed on controlled release drug delivery systems to maintain the therapeutic efficacy of these drugs. There are a number of mechanisms that can provide such controlled release of drugs, including transdermal patches, implants, bioadhesive systems, and microencapsulation [7]. Newer drug delivery technologies currently attracting attention include inhaled and sustained release injectable peptide / protein drugs from biodegradable polymers. Nevertheless, the ideal drug delivery system has not yet been attained. Therefore, interdisciplinary approaches are being sought. These alternative approaches should, at least conceptually, address those

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unsolved issues that still engage the drug delivery scientist. These include drug targeting (of peptides, proteins, and DNA) to improve cancer chemotherapy and cardiovascular treatment, improved control of the rate of release from implanted dosage forms, pulsatile drug delivery, closed-loop systems that can sense a change in the bodies biochemistry leading to modulated rate or extent of drug delivered, and the ability to implant xenogeneic cells in humans to more naturally control a range of diseases, including diabetes and neurodegenerative conditions.

3. Microfabrication technology The use of traditional microfabrication techniques, the same processing techniques used to manufacture microelectronic chips, is a recent and alternative method of creating drug delivery platforms. Microelectronic process engineering was a discipline that developed due to the rapid growth of the integrated circuit industry. Traditionally, microelectromechanical systems (MEMS) research has been used to produce functional devices on the micron scale, such as sensors, switches, filters, and gears, from silicon, the dominant material used throughout the IC industry. These devices are fabricated by the repeated application of unit process steps such as thin-film deposition, photolithography, and etching. Such microfabrication techniques allow for the precise control over surface microarchitecture, topography, and feature size. Although research on microfabricated devices for biomedical applications (BioMEMS) has rapidly expanded in recent years, relatively few researchers have concentrated on therapeutic applications of microfabrication technology such as drug delivery. The use of microtechnology to tailor the size, shape, reservoir number, reservoir volume, unidirectional openings and surface characteristics of the drug delivery vehicle in conjunction with appropriate surface chemistry are potentially influential in the area of controlled release.

4. Microneedles for transdermal drug delivery One alternative to oral delivery and intravenous

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injection is the administration of drugs across the skin. This approach seeks to avoid any degradation of the molecules in the gastrointestinal tract and first-pass effects of the liver associated with oral drug delivery as well as the pain of intravenous injection [10–14]. It also offers the possibility to continuously control the delivery rate over extended periods of time [14]. However, conventional transdermal drug delivery is severely hindered by the outer 10–20 mm of skin, a barrier of dead tissue called the stratum corneum [15]. The development of microneedles for transdermal drug delivery came about as an approach to enhance the poor permeability of the skin by creating microscale conduits for transport across the stratum corneum [14]. The development of microneedles that are long and robust enough to penetrate this layer of skin, but short enough to avoid stimulating nerves has the potential to make transdermal delivery of drugs more effective [15].

4.1. Microfabrication of silicon microneedle arrays By adapting microfabrication technology, threedimensional arrays of sharp-tipped microneedles can be made for transdermal drug delivery [10,14,15]. To fabricate microneedles, a deep reactive ion etching process is commonly used. In this process, a chromium masking material is deposited onto silicon wafers and patterned into dots that have a diameter approximately equal to that of the base of the desired needles. When placed in the reactive ion etcher, the wafers are exposed to carefully controlled plasma of fluorine and oxygen, which causes a deep vertical, etch and slight lateral underetching. The regions on the wafer that are protected by chromium remain and eventually form the microneedles. Etching is allowed to proceed until the masks are undercut and fall off, leaving behind an array of silicon spikes [14]. The aspect ratio of the microneedles can be adjusted by simply modifying the ratio of flow rates of SF 6 and O 2 . Hollow silicon needles can also be fabricated using deep reactive ion etching in an inductively coupled reactive ion etcher. The deep etch creates arrays of holes through the silicon wafer (the needle lumen) and the microneedles are formed by reactive ion etching around these holes [15]. Arrays of solid silicon microneedles have been fabricated with individual needles measuring 150 mm

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in length, 80 mm in diameter at the base, with a radius of curvature less than 1 mm [14]. Hollow needles have also been microfabricated with similar dimensions, but containing hollow bores anywhere from 5 to 70 mm in diameter, depending on design [15]. To test their durability, the solid needles were inserted into skin with gentle pushing, an approximate force of 10 N. All but a few percent of the microneedles remained intact. Even in these few needles, only the top 5–10 mm was damaged [14,15]. Additionally the array of microneedles could also be removed without additional damage and could also be reinserted into skin multiple times.

4.2. Transdermal transport studies Quantification of transdermal transport of various molecules with and without inserted microneedle arrays was used to assess any increase in skin permeability leaving (Fig. 1). Insertion of the microneedles increased permeability only 1000-fold because the microneedles or the silicon plate may have blocked access to the microscopic holes [14]. When the microneedles were removed after 10 s, permeability increased by 10 000-fold [14]. Removal after 1 h increased skin permeability by 25 000-fold [14]. Elevated permeability after microneedle insertion was found to remain at approximately the same

level for as long as 5 h [14]. Hollow microneedles were also capable of insertion into the skin without any extensive damage to the microneedles or skin [15]. In addition, the improved design of these needles increased skin permeability further still [15].

4.3. Other microneedles More recently, the original microneedle design has been further refined to provide better control over drug delivery. Silicon microhypodermic needles have been fabricated in combination with heat-controlled bubble pumps [15]. Hollow metal microneedles have been fabricated by defining molds in epoxy and filling them by electrodepositing metal [15]. Similarly, polysilicon microneedles have been fabricated with reusable molds [15,16]. Polysilicon microneedles are likely to be most cost effective and have the potential to produce single use disposable platforms [15]. These types of needles, combined with a pressurized reservoir to generate a drug delivery pump, have already been incorporated into a wearable drug infusion system to deliver insulin [17]. Furthermore, by modifying needle dimension and design to incorporate multiple channels and ports, optimized microhypodermic needles and microprobes can be developed for cellular, local tissue, or systemic delivery [15].

Fig. 1. (A) Scanning electron micrograph of microneedles made by reactive ion etching technique. (B) Microneedle tips inserted across epidermis. The underside of the epidermis is shown, indicating that the microneedles penetrated across the tissue and that the tips were not damaged. Arrows indicate some of the microneedle tips. Reproduced with permission from Ref. [14].

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5. Implanted microchip for localized drug delivery Microfabrication technology has also created a new class of controlled release systems for drug delivery based on programmable devices. These devices are particularly intriguing due to their small size, potential for integration with microelectronics and their ability to store and release chemicals on demand [18]. With the recent advancements in biosensors and micromachining, implanted responsive drug release systems are becoming more plausible.

5.1. Microchip design The first experimental demonstration of a microchip with potential application in drug delivery was described in Nature [19]. The ultimate goal was to develop a microfabricated device devoid of moving parts, but with the ability to store and release multiple chemical substances. The device was fabricated by the sequential processing of a silicon wafer using microelectronic processing techniques including UV photolithography, chemical vapor deposition, electron beam evaporation and reactive ion etching [19]. The experimental prototype was a 17 mm 3 17 mm 3 310 mm square silicon device containing an array of 34 square pyramidal reservoirs etched completely through the wafer [18,19].

5.2. Irreversible metallic valves The 25-nl reservoirs were sealed at one end by a thin membrane of gold to serve as an anode in an electrochemical reaction [18,19]. One other electrode was placed on the device to serve as a cathode. The reservoirs were filled through the open end with the chemical to be released by either microsyringe pumps or inkjet printing in conjunction with a computer-controlled alignment apparatus [18,19]. The open ends of the reservoirs were then covered with a thin adhesive plastic and sealed with waterproof epoxy [18,19]. When submerged in an electrolyte, ions form a soluble complex with the anode material in its ionic form [18]. An applied electric potential oxidizes the anode membrane, forming a soluble complex with the electrolyte ions [18]. The

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complex dissolves in the electrolyte, the membrane disappears, and the chemical is released and allowed to diffuse form the reservoir. The time at which release occurs from each individual reservoir is determined by the time at which the reservoir’s anode membrane is removed [18]. Each reservoir, or a group of reservoirs, may be independently addressed by demultiplexing [18]. This allows each anode to have its own conducting path and electric potential can be applied to any given combination of reservoirs at any given time [18,19]. However, the rate of release from the reservoir is a function of the dissolution rate of the materials in the reservoir and the diffusion rate of these materials out of the reservoir [18]. Therefore, the rate of release from an individual reservoir can be controlled by proper selection of the materials (e.g. pure drugs, or drugs with polymers) placed inside the reservoir [18]. Using a material that quickly dissolves once the reservoir is opened can be used to achieve pulsatile release whereas a material that dissolves slowly after the reservoir is opened can be used to achieve sustained release (Fig. 2) [18].

5.3. Reversible polymeric valves An alternative to the use of irreversible metallic valves is a microchip using reversible polymeric valves. The use of ‘artificial muscle’ valves in conjunction with silicon micromachined drug release structures can render a microchip responsive to a patient’s therapeutic requirements and deliver certain amount of a drug in response to a biological stimulus [20]. ‘Artificial muscle’ refers to a chemomechanical actuator consisting of a blend of a hydrogel and an electronically conducting redox polymer [21]. The redox polymer is sensitive to pH, applied potential, and the chemical potential of its microenvironment whereas the hydrogel provides a cross-linked network of hydrophilic homo / copolymers that exhibit dramatic swelling and shrinking upon changes in pH, solvent, temperature, electric field, or ambient light conditions [21]. By electropolymerizing these polymers onto electrodes, reservoirs can be opened or closed, and the drug compound released or retained, via the swelling and shrinking processes of the polymer system in response to electrochemical actuation (Fig. 3) [20].

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Fig. 2. Photographs of a prototype microchip: the electrode-containing front side and the back side with openings for filling the reservoirs. Scale bar 10 mm. Reproduced with permission from Ref. [18].

5.4. Current developments Current developments on microchip systems consist of integrating active components. Combining battery clocks, reference electrodes and biosensors with the microchip could provide a single package for implantation. Another logical extension of the controlled-release microchip would be the development of a passive, polymer microchip that contains no electronics, power sources, or microprocessors [18]. Each of the reservoirs would be covered by a cap of degradable material, or a nondegradable material of known permeability for the drug, or left uncapped. Time and rate of release from the reservoir would then be dictated by the degradation rate of the cap or diffusion of the drug [18]. This type of polymeric microchip device would have the additional advantage of being biodegradable, and once

implanted for drug delivery applications need not be removed.

6. Bioadhesive microparticles for oral drug delivery Oral drug delivery is one of the most preferred methods of drug administration due to its non-invasive nature. However, it is generally not a viable method for peptide and protein delivery. The human GI tract resists absorption of peptides, proteins, and other large molecules until they are broken down into smaller molecules. The acidic environment of the stomach combined with an array of enzymes and physical barriers in the intestines either destroy or prevent absorption of nearly all macromolecules.

Fig. 3. (A) A Ag /AgCl and IrO x valve electrode in the same micromachined drug delivery cavity. Both electrodes are 30 3 30 mm. (B) SEM micrograph of ‘artificial muscle’ grown on TEM gold grid coated with poly-HEMA in holes (38.5 mm 3 38.5 mm) of a drug reservoir. Reproduced with permission from Ref. [20].

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This has led to the development of oral delivery systems that can potentially enhance the delivery of peptides utilizing mechanisms such as use of protective coatings [22], targeted delivery [23], permeation enhancers [24], and protease inhibitors [25]. Bioadhesive drug delivery systems have also generated considerable interest due to their potential for prolonging the residence time at the site of drug action or adsorption. Localization of the delivery system at a given target site would intensify its contact with the mucosal epithelial barrier, thereby increasing the drug concentration gradient due to intense contact [26,27]. Rather than having an implanted controlled-release microchip for local drug delivery, such systems could also be fabricated for specific targeting. By developing inorganic or polymeric reservoir-containing particles on the micron scale, and grafting bioadhesive agents on their releasing side, these particles could be adapted for use as a bioadhesive controlled release oral drug delivery system. This type of system could improve the effectiveness of treatment by first, targeting and localizing a drug at a specific site, inhibiting dilution of the drug in body fluids, thereby helping to maintain the drug concentration at the optimal concentration between effective and toxic levels. Micromachined platforms, when combined with complementary approaches, may address some of the shortcomings of current oral delivery systems for peptides and proteins by combining several features into a single drug delivery platform. First, one can achieve control over the size and shape of the delivery device. Unlike other spherical drug delivery particles, microfabricated devices may be designed to be flat, thin, and disc-shaped to maximize contact area with the intestinal lining and minimize the side areas exposed to the constant flow of liquids through the intestines. The size of the particles can be selected to be small enough to have good contact with the undulations of the intestinal wall and large enough to avoid endocytosis of the entire particle. While endocytosis of nanoparticles has been proposed as a method to enhance transport of large molecules across the intestinal barrier, this process can destroy the macromolecule. Secondly, one can selectively attach bioadhesive agents onto the device surface using relatively simple surface chemical modification strategies. Finally, micromachining pre-

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sents the opportunity to create multiple reservoirs of desired size to contain not just one, but many drugs / biomolecules of interest [28].

6.1. Silicon dioxide microparticles Microparticles have been successfully fabricated from silicon dioxide using microfabrication protocols [28]. These particles are then easily modified by silane chemistry to introduce to sites for the attachment of biological molecules.

6.1.1. Fabrication of microparticle body To fabricate the particle body an etch stop layer was created by growing a thermal oxide under wet conditions. Low-pressure chemical vapor deposition was then used to deposit a sacrificial layer of polycrystalline silicon atop the thermal oxide by lowpressure chemical vapor deposition. Next, a layer of low temperature silicon dioxide (LTO) was deposited to form the device layer. Negative lithography was carried out to mask define the particle shape. A reactive ion etch (RIE) with SF 6 and O 2 was used to define the actual LTO particles and any remaining photoresist was then removed in negative photoresist remover. 6.1.2. Fabrication of microparticle reservoir Positive lithography was carried out using infrared (IR) backside alignment to define the wells. The wafers were then time-etched in buffered oxide etchant to carve out the wells. Any remaining photoresist was removed. The welled microdevices were then released into solution by etching the sacrificial polysilicon layer with KOH. The KOH solution was diluted with deionized water and filtered to isolate the micro devices. As seen in Fig. 4A, the particles were fabricated with well-defined features across the entire wafer. Released microparticles are shown in Fig. 4B. The particles were uniform and semi-transparent due to their polycrystalline nature. 6.1.3. Surface modification chemistry Chemistries used to link biologically active molecules, such as proteins, on to the releasing side of the silicon dioxide microparticles have also been developed using traditional silane chemistry and car-

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Fig. 4. (A) An array of welled SiO 2 microparticles. (B) Released microparticles.

bodiimide coupling reagents. First, amine groups are formed on the surface of unreleased microparticles through hydroxylation of the silicon, followed by silanization in 2% v / v APTES at room temperature for at least 1.5 h. Biological molecules that contain carboxyl groups were coupled with carbodiimide and N-hydroxysuccinimide to form a reactive intermediate ester that is susceptible to attack by amines. This reaction can be applied to self-assemble a monolayer of covalently coupled biological molecules to the surface of the microparticles.

6.2. Poly(methyl methacrylate) microparticles Such a system has also been developed based on the polymer polymethylmethacrylate (PMMA). PMMA is an ideal material for such a MEMS-based bioadhesive system because it is biocompatible and already used in many biomedical applications, is commonly used as a resist in photolithographic applications, and contains a functional methyl ester group for potential surface modification. The current prototype is a square particle 150 mm across and 3 mm thick containing square reservoirs 80 mm across and 2 mm deep.

6.2.1. Fabrication of microparticle body To fabricate these particles, PMMA was spun on to clean silicon wafers. Positive lithography was used to expose the area between the particles and isolate the particle bodies. The unmasked area was then etched completely through in the reactive ion etcher using an O 2 plasma and any remaining resist was removed. The thickness of the particles can be

modified by changing the spin rate or by spinning on additional layers of PMMA. Additionally, altering the masked area of the wafer easily changes the shape and surface area of the particles.

6.2.2. Fabrication of microparticle reservoir A second positive lithography was carried out to expose the intended reservoir areas in the PMMA. This area was then etched 1–2 mm deep using an oxygen plasma and any remaining photoresist was removed. The dimensions of the reservoir can be altered by changing the masked area and their depth can be modified by changing the time and / or flow rate of plasma in the RIE (Fig. 5). By creating smaller wells, a series of multiple wells can be etched into the particles to create separate reservoirs for a combination of drugs or permeation enhancers. Since the PMMA is adherent to the surface of silicon by linkage to the native oxide layer, the wafer was soaked in basic solution to break this bonding and immediately release the particles. 6.2.3. Surface modification chemistry Heterogeneous modification of PMMA with Nlithioethylenediamine as the aminolyzing agent leads to amination of its ester groups [29,30]. This reaction layer of amine sites tethered to the PMMA backbone by stable amide groups. Amine groups were placed on the surface of the releasing side of PMMA microparticles by reacting unreleased microparticles with N-lithioethylenediamine for 10 min. Biological molecules were linked to these amine sites using the same traditional carbodiimide coupling reagents used for silicon dioxide microparticles.

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Fig. 5. Arrays of 150 mm 3 150 mm PMMA particles with (A) 50 mm 3 50 mm, (B) 80 mm 3 80 mm, (C) 100 mm 3 100 mm, and (D) multiple 28 mm 3 28 mm wells.

6.3. Release mechanism These reservoirs can be filled with pico- to nanoliters of a polymeric solution with microinjectors. Water quickly evaporates from these reservoirs leaving behind the drug contained in polymer which acts as a timed-release plug. Using a specific type of polymer predetermines the time and rate of release of drug from the reservoir; for example, a hydrogel that swells in response to a specific pH, solvent or temperature or a polymer with a known dissolution rate. Different polymers with various dissolution rates can be used in each reservoir to obtain controlled release of separate compounds.

6.4. Microparticle bioadhesion to intestinal epithelium The current microparticle model is designed to specifically target the intestinal epithelium. The microparticles were first rendered bioadhesive by forming avidin-coupled surfaces using the chemistry previously described. Commercially available biotinylated lectins were then attached to the microparticle surface by taking advantage of the strong interaction and affinity of avidin and biotin in nature.

6.4.1. Lectins Lectins are a class of carbohydrate binding proteins or glycoproteins of non-immune origin. They have been deemed a ‘second generation’ bioadhesive because they bind to cell-surface glycoconjugates in a complementary way analogous to ligand–receptor interactions. Tomato lectin is of special interest in intestinal targeting due to its stability in low pH environments and its non-toxicity. Furthermore, it has been shown to bind selectively to the small intestine epithelium [31], resist digestion in the alimentary canal of rats and bind to rat intestinal villi without disruption of the lectin integrity [32].

6.4.2. Caco-2 binding In vitro studies using a Caco-2 model of the intestinal epithelium were used to demonstrate the specific binding of the tomato lectin-conjugated PMMA particles. The tomato lectin-conjugated particles showed, on average, a marked increase in binding almost five times greater than the binding of unmodified particles over an entire period of 2 h (Fig. 6). Furthermore, the modified particles once bound, remained bound whereas only 20% of originally bound unmodified particles remained bound after 20 min. Currently, in vivo studies are being

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The disease manifests itself as hyperglycemia. Insulin remains the mainstay of virtually all type 1 DM and many type 2 DM patients and in most cases is administered subcutaneously. However, the kinetics of insulin administered by this route do not mimic the normal rapid rise and decline of insulin secretion in response to ingested nutrients. Efforts to address the short-comings of current subcutaneous administration of insulin, including the use of complex multidose regimens, have led to the development of other dosage forms and routes of administration such as ‘needleless’ injectors, constant infusion pumps, and inhaled insulin. These newer approaches still suffer from the same general issue plaguing current subcutaneous administration. Fig. 6. Binding of unmodified and lectin-conjugated PMMA microparticles to Caco-2 monolayers.

conducted to determine the bioadhesive properties of these particles in the gastrointestinal tract of rats.

6.5. Future developments By replacing the molecule attached to the microparticles, an array of cells and tissues can be targeted. For example, if the lectin is substituted with an antibody that selectively binds to tumor cells in the colon, the microparticles could actively seek out cancerous masses in the colon and deliver anticancer drugs directly. This would allow the high concentration of drug to be locally delivered while keeping the systemic concentration at a low level. Moreover, microfabricated polymer particles with the same targeting abilities, but small enough for injection, could be developed for direct delivery into the circulatory system.

7. Nanoporous immunoisolating biocapsules Diabetes mellitus (DM) represents a serious medical problem. In the US alone, it is the third leading cause of death. While the majority of patients have type 2 diabetes, about 10% of all patients diagnosed with DM are insulin-dependent (type 1). In both cases, disease is caused by decreased circulating concentrations of insulin and decreased response of peripheral tissue to insulin (insulin resistance) [33].

7.1. Allotransplantation A potentially useful approach, which has proven effective in only a handful of cases, is the allotransplantation of islets or whole pancreases from a suitable human donor into a diabetic recipient. Researchers in Canada recently reported successful transplantation of islet cells in type 1 DM patients [34]. Although the potential complications of immunosuppressive therapy were reduced by avoiding the use of glucocorticoids, each transplant required two harvestings of islet cells from organ donors. Moreover recipients are still required to take immune suppressing drugs for the rest of their lives. These immunosuppressive drugs are toxic and have potential adverse side effects, including cancer. For this reason, an islet or pancreas allotransplant is normally carried out only in conjunction with a kidney transplant, for which immunosuppression is required in any case. Because of the toxicity of immune suppressing drugs, and the shortage of organ donors, islet and pancreas allotransplantation appears to hold limited promise as a cure for diabetes. A method then is required to sequester the islets from the body’s immune system which is able to recognize and reject these xenogeneic cell grafts. For the past 20 years, investigators have focused on a range of microencapsulation methods most commonly involving sodium alginate and another polycationic substance such as polylysine. These materials have been used in an attempt to create a semipermeable membrane

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capable of blocking immune molecules such as IgG, cytokines, and cell-secreted antigens from reaching the encapsulated xenogeneic islet cells while allowing glucose and insulin to freely diffuse through the barrier [35]. However, this approach has proven generally unsuccessful due to mechanical rupture of the membrane, biochemical instability, incompatibility with islet cell heterogeneity, and broad pore size distributions [35–39]. When the barrier between the xenogeneic cells and the external bioenvironment is compromised, these foreign cells are subject to various endogenous cells and antibodies as well as complement and a host of cytokines such as tumor necrosis factor, all of which can inflict cell damage. As a result, the use of polymeric microcapsules for allotransplantation has been unsuccessful clinically in the absence of immunosuppression [35,37,39].

7.2. Biocapsule design Microfabrication techniques have been applied to create a biocapsule for effective immunoisolation of transplanted islet cells for treatment of diabetes [40]. The fabrication of nanochannels in the membrane structure consists of two steps: (1) surface micromachining nanochannels in a thin film on the top of a silicon wafer, and (2) releasing the membrane by etching away the bulk of the silicon wafer underneath the membrane. These nanopore membranes (Fig. 7) are designed to allow the permeability of glucose, insulin, and other metabolically active products, while at the same time, preventing the passage of cytotoxic cells, macrophages, antibodies, and complement. The membranes are bonded to a capsule that houses the pancreatic islet cells. Because the difference in the size of insulin, which must be able to pass freely through the pores and the size of IgG immunoglobulins, which must be excluded, is only a matter of a few nanometers, the highly uniform pore size distribution provided by micromachine membranes is essential for effective immunoisolation and therapeutic effect. Control of pore sizes in the tens of nanometers has recently been suggested as probably the most realistic way to achieve immunoisolation [41,42]. The use of unconventional biomaterials such as silicon and silicon dioxide provides a means to encapsulate pancreatic islet cells in devices that are thermally,

Fig. 7. Micrograph of a biocapsule membrane with 24.5-nm pores.

chemically, and mechanically stable and retrievable. It is also expected that improved dynamic response of islets can be obtained due to the limited membrane thickness (Fig. 8) compared with the thickness of conventional polymeric membranes prepared from alginate and polylysine (100–200 mm) [41]. It is crucial that rapid secretion kinetics, particularly in

Fig. 8. Insulin secretory profile through differing pore sizes.

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the first phase of insulin release, be maintained over time to provide physiologic feedback control of blood glucose concentrations. Another key feature of an implanted biocapsule system is the role of neovascularization. The membranes proposed for testing have outer openings of 2 by 2 mm while the inner diffusion channels have a pore size of between 10 and 30 nanometers. Studies have shown that neovascularization at the membrane–tissue interface occurs in membranes having pore sizes large enough to allow complete penetration by host cells (0.8–8 mm) [41]. Thus, it is expected that neovascularization can occur at the large openings while not penetrating into the nanometer pores. This phenomenon has two key advantages: (1) the ability to rapidly deliver insulin into the blood stream through new blood vessel growth while (2) limiting pore clogging or fouling.

7.3. Biocompatibility and cytotoxicity studies Preliminary biocompatibility and cytotoxicity tests have been carried out by examining the cell morphology, growth, and function of test cell lines placed in contact with arrays of membranes, with promising preliminary results. The biocompatibility was evaluated via direct contact tests by cultivating several different cell lines such as macrophages, fibroblasts, and HeLa cells, as well as isolated primary islets of Langerhans both on the wafer surface and within the porous wafer pockets. All cells were seeded on silicon culture wafers, observed via light microscope, stained for cell viability and functionality, and counted with a hemocytometer. All cell types had normal growth characteristics, morphology, and greater than 90% viability [43,44]. Results indicate that the insulin secretion by encapsulated islets and subsequent diffusion through the biocapsule membrane channels is similar to that of unencapsulated islets for both 3 mm and 78 nm pore sized membranes, with insulin diffusion though the membrane occurring within 10 min of stimulation. Fig. 8 shows the typical insulin release profile in response to stimulatory (16.7 mM) glucose medium over 1 h under static incubation for 78-, 66and 18-nm pore-sized membranes. This profile indicates that insulin and glucose diffusion occur at sufficiently high rates through the microfabricated

Fig. 9. IgG diffusion through microfabricated biocapsules of three different pore sizes.

membrane to ensure nutrient exchange for encapsulated islet cells. These experiments show that no diffusion barrier is formed by the membrane for glucose and insulin, while taking into account the effect of rotation on mass transfer. In addition, microfabricated biocapsule membranes can be tailor made to attain desired IgG diffusion kinetics (Fig. 9). At the same time, the deselection of IgG requires absolute pore dimensions below 18 nm. It is noted that the percent of IgG diffusion (concentration of IgG that passes through the membrane) was less than 0.4% after 24 h and 2% after over 150 h through the 18-nm membranes [44–46]. It may be possible to design nanopore membranes which achieve a more constant rate of drug delivery, avoiding the ‘burst effect’. By precisely controlling pore size, pore length and pore density, the nanopore membrane fitted into a polymeric capsule suitable for subcutaneous implantation can serve as a diffusion barrier for a variety of biological drugs.

8. Conclusions The race to find effective diagnostic and therapeutic tools is under way, as scientific and engineering disciplines uncover and elucidate more about the human pathologic condition than ever before. Although we are getting closer to the clinical application of intelligent drug delivery devices, many challenges remain for the future. The convergence of microtechnology and biology will lead to new approaches in drug delivery and may provide advan-

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tages over existing technologies. By focusing efforts at the microscale, we have the unique ability to engineer control over the cellular environment, leading to novel ways in which we can control molecular delivery and cell / tissue interactions. The future challenge lies in assembling and applying our collective knowledge to develop functional and clinically relevant therapeutic delivery devices.

Acknowledgements Funding is gratefully acknowledged from The Whitaker Foundation, NSF ECS9820829, NSF Career, and iMEDD, Inc. Also, special thanks to those who have contributed to this work: Lara Leoni, Mike Lubeley, Chris Bonner, and Aamer Ahmed of UIC; colleagues from iMEDD, Inc.; and Professors Derek Hansford and Mauro Ferrari from OSU.

References [1] P.L. Gourley, Semiconductor microlasers: a new approach to cell-structure analysis, Nat. Med. 2 (1996) 942–944. [2] J. Drews, Drug discovery: a historical perspective, Science 287 (2000) 1960–1964. [3] D.J. Anderson, K. Najafi, S.J. Tanghe, D.A. Evans, K.L. Levy, J.F. Hetre, X. Xue, J.J. Zappia, K.D. Wise, Batchfabricated thin-film electrodes for stimulation of the central auditory system, IEEE Trans. Biomed. Eng. 36 (1989) 693– 704. [4] J. Evans, D. Liepmann, A.P. Pisano, Planar laminar mixer, Proc. IEEE MEMS Workshop 10 (1997) 96–101. [5] J. Deutsch, D. Motlagh, B. Russell, T.A. Desai, Fabrication of microtextured membranes for cardiac myocyte attachment and orientation, J. Biomed. Mater. 53 (2000) 267–275. [6] M.L. Reed, C. Wu, J. Kneller, S. Watkins, D.A. Vorp, A. Nadeem, L.E. Weiss, K. Rebello, M. Mescher, A.J.C. Smith, W. Rosenblum, M.D. Feldman, Micromechanical devices for intravascular drug delivery, J. Pharm. Sci. 87 (1998) 1387– 1394. [7] D.D. Breimer, Future challenges for drug delivery, J. Controlled Release 62 (1999) 3–6. [8] J.T. Santini, A.C. Richards, R.A. Scheidt, M.J. Cima, R.S. Langer, Microchip technology in drug delivery, Ann. Med. 32 (2000) 377–379. [9] S.S. Davis, L. Illum, Drug delivery systems for challenging molecules, Int. J. Pharm. 176 (1998) 1–8. [10] D.L. Polla, A.G. Erdman, W.P. Robbins, D.T. Markus, J. Diaz-Diaz, R. Rizq, Y. Nam, H.T. Brickner, Microdevices in medicine, Annu. Rev. Biomed. Eng. 2 (2000) 551–576.

327

[11] J. Hadgraft, R.H. Guy (Eds.), Transdermal Drug Delivery: Developmental Issues and Research Initiatives, Marcel Dekker, 1989. [12] E.W. Smith, H.I. Maibach (Eds.), Precutaneous Penetration Enhancers, CRC Press, 1995. [13] B.G. Amsden, M.F.A. Goosen, Transdermal delivery of peptide and protein drugs: an overview, AIChE J. 41 (1995) 1972–1997. [14] S. Henry, D.V. McAllister, M.G. Allen, M.R. Prausnitz, Microfabricated microneedles: a novel approach to transdermal drug delivery, J. Pharm. Sci. 87 (1998) 922–925. [15] D.V. McAllister, M.G. Allen, M.R. Prausnitz, Microfabricated microneedles for gene and drug delivery, Annu. Rev. Biomed. 2 (2000) 289–313. [16] D. Lieppmann, A.P. Pisano, B. Sage, Microelectromechanical systems technology to deliver insulin, Diabetes Technol. Ther. 1 (1999) 469–476. [17] J.D. Zhan, A.A. Peshmukh, A.P. Pisano, D. Liepmann, Continuous on-chip micropumping through a microneedle, in: Int. 14th Conf. MEMS, 2001, pp. 503–506. [18] J.T. Santini, A.C. Richards, R. Scheidt, M.J. Cima, R. Langer, Microchips as controlled drug-delivery devices, Angew. Chem. Int. Ed. Engl. 39 (2000) 2396–2407. [19] J.T. Santini, M.J. Cima, R. Langer, A controlled-release microchip, Nature 397 (1999) 335–338. [20] L. Low, S. Seetharaman, K. He, M.J. Madou, Microactuators toward microvalves for responsive controlled drug delivery, Sensors Actuators B Chem. 67 (2000) 149–160. [21] M. Madou, J. Florkey, From batch to continuous manufacturing of microbiomedical devices, Chem. Rev. 100 (2000) 2679–2692. [22] M. Saffran, G.S. Kumar, D.C. Neckers, J. Pena, R.H. Jones, J.B. Field, Biodegradable azopolymer coating for oral delivery of peptide drugs, Biochem. Soc. Trans. 18 (1990) 752– 754. [23] J.W. Fara, R.E. Myrback, D.R. Swanson, Evaluation of oxprenolol and metoprolol Oros systems in the dog: comparison of in vivo and in vitro drug release, and of drug absorption from duodenal and colonic infusion sites, Br. J. Clin. Pharmacol. 19 (1985) 91S–95S. [24] A. Fasano, S. Uzzau, Modulation of intestinal tight junctions by Zonula occludens toxin permits enteral administration of insulin and other macromolecules in an animal model, J. Clin. Invest. 99 (1997) 1158–1164. [25] U.I. Schwarz, T. Gramatte, J. Krappweis, R. Oertel, W. Kirch, P-glycoprotein inhibitor erythromycin increases oral bioavailability of talinolol in humans, Int. J. Clin. Pharmacol. Ther. 38 (2000) 161–167. [26] C.M. Lehr, Lectin-mediated drug delivery: the second generation of bioadhesives, J. Controlled Release 65 (2000) 19–29. [27] G. Poncel, J. Irache, Specific and non-specific bioadhesive particulate systems for oral delivery to the gastrointestinal tract, Adv. Drug Deliv. Rev. 34 (1998) 191–219. [28] A. Ahmed, C. Bonner, T.A. Desai, Bioadhesive microdevices with multiple reservoirs: a new platform for oral drug delivery, J. Controlled Release 81 (2002) 291–306.

328

S.L. Tao, T. A. Desai / Advanced Drug Delivery Reviews 55 (2003) 315–328

[29] B. Karandikar, J. Puschett, K. Matyjaszewski, Homogeneous and heterogeneous modification of poly(methyl methacrylate) with ethylene diamine, Polym. Prep. Am. Chem. Soc. Div. Polym. Chem. 30 (1989) 250–251. [30] A.C. Henry, T.J. Tutt, M. Galloway, Y. Davidson, C.S. McWhorter, S. Soper, R. McCarley, Surface modification of poly(methyl methacrylate) used in the fabrication of microanalytical devices, Anal. Chem. 72 (2000) 5331–5337. [31] B. Carreno-Gomez, J.F. Woodley, A.T. Florence, Studies on the uptake of tomato lectin nanoparticles in everted gut sacs, Int. J. Pharm. 183 (1999) 7–11. [32] D. Kilpatrick, A. Pusztai, G. Grant, C. Graham, S. Ewen, Tomato lectin resists digestion in the mammalian alimentary canal and binds to intestinal villi without deleterious effects, FEBS Lett. 185 (1985) 299–305. [33] S.N. Davies, D.K. Granner, The Pharmacological Basis of Therapeutics, McGraw-Hill, 1996. [34] S.S. Davies, Biomedical applications of nanotechnology— implications for drug targeting and gene therapy, Trends Biotechnol. 15 (1997) 217–224. [35] R.P. Lanza, D.K. Cooper, Xenotransplantation and cell therapy: progress and controversy, Mol. Med. Today 5 (1999) 105–106. [36] C.K. Colton, E.S. Avgoustiniatos, Bioengineering in development of the hybrid artificial pancreas, ASME J. Biomech. Eng. 113 (1991) 152–170. [37] P. Lacy, O.D. Hegre, A. Gerasimidi-Vazeou, F.T. Gentile, K.E. Dionne, Maintenance of normoglycemia in diabetic mice by subcutaneous xenograft of encapsulated islets, Science 254 (1991) 1728–1784. [38] P. Soon-Shiong, M. Otterie, G. Skjak-Braek, O. Smidsrod, R. Heintx, R.P. Lanza, T. Espevik, An immunologic basis for

[39] [40]

[41]

[42]

[43]

[44]

[45]

[46]

the fibrotic reaction to implanted microcapsules, Transplant. Proc. 23 (1991) 758–759. R.P. Lanza, J.L. Hayes, W.L. Chick, Encapsulated cell technology, Nat. Biotechnol. 14 (1996) 1107–1111. T.A. Desai, W.H. Chu, J.K. Tu, G.M. Beattie, A. Hayek, M. Ferrari, Microfabricated immunoisolating biocapsules, Biotechnol. Bioeng. 57 (1998) 118–120. M. Brissova, I. Lacik, A.C. Powers, A.V. Anilkumar, T. Wang, Control and measurement of permeability for design of microcapsule cell delivery system, J. Biomed. Mater. Res. 39 (1998) 61–70. T. Wang, I. Lacik, M. Brissova, A.V. Anilkumar, A. Prokop, D. Hunkeler, R. Green, K. Shahrokhi, A.C. Powers, An encapsulation system for the immunoisolation of pancreatic islets, Nat. Biotechnol. 15 (1997) 358–362. T.A. Desai, D. Hansford, M. Ferrari, Characterization of micromachines membranes for immunoisolation and bioseparation applications, J. Biomed. Microdev. 4132 (1999) 1–11. T.A. Desai, W.H. Chu, G. Rasi, P.S. Vallebona, E. Guarino, M. Ferrari, Microfabricated biocapsules provide short-term immunoisolation of insulinoma xenografts, Biomed. Microdev. 1 (1999) 131–181. T.A. Desai, D.J. Hansford, L. Leoni, M. Essenpreis, M. Ferrari, Nanoporous anti-fouling silicon membranes for implantable biosensor applications, Biosens. Bioelectron. 15 (2000) 453–462. L. Leoni, T. Boriarski, T.A. Desai, Characterization of nanoporous membranes for immunoisolation: diffusion properties and tissue effects, J. Biomed. Microdev. 4 (2002) 131–139.