CONTENTS
9.1.1 Anatomy and Physiology of the Skin: Implications for Transdermal Drug Delivery
9.1.2 Advantages of the Transdermal Route for Drug Delivery
9.1.3 Limiting Factors for Transdermal Drug Delivery
9.2 Mechanical and Biological Properties of the Principal Skin Layers
9.2.1 Stratum Corneum: The Key Mechanical Barrier
9.2.2 Viable Epidermis and Dermis
9.3 Approaches to Transdermal Drug Delivery
9.3.1 Passive Approaches Relying on Diffusion
9.3.1.1 Transdermal Therapeutic Systems
9.3.2 Active Approaches Requiring Application of Energy
9.3.3 Physical Approaches That Breach the Stratum Corneum
9.3.3.1 Intradermal Injections
9.3.3.4 Microneedles, Microprojections, and Arrays Thereof
9.3.3.5 Nanopatch™: A Case Study
Transdermal drug delivery is an emerging field with key applications in local delivery of molecules to local tissue sites, through to systemic delivery applications under sustained-release conditions. Historically, transdermal delivery has been limited to molecules that fit a narrow physicochemical profile (low molecular weight, adequate solubility in both oil and water, high partition coefficient); however, the range of deliverable drugs is rapidly expanding on all fronts, thanks to advances primarily in (1) enhancing the permeability of the outermost skin layer (stratum corneum [SC]), (2) increasing the driving force for drug transport across the SC, (3) physical approaches that bypass the SC altogether, and (4) novel combinations of these methods. With ≈20 small molecule drug delivery systems (DDS) currently FDA-approved involving transdermal approaches, based on current trials in small animals and humans, it is likely that delivery of genetic and cellular therapies, vaccines, therapeutic proteins, and nanoparticle-encapsulated systems will be available in the near future.
9.1.1 ANATOMY AND PHYSIOLOGY OF THE SKIN: IMPLICATIONS FOR TRANSDERMAL DRUG DELIVERY
The skin is the largest organ of the human body (1.5–2.5 m2 surface area); its structure is described in Chapter 4 (Section 4.5.1 and Figure 4.6). It is composed of three principal skin layers: the SC, 10–20 μm; the viable epidermis (VE), 50–100 μm; and the dermis, 1–2 mm. The dermis is highly vascularized, while the SC/VE is not, yet capillary loops in the reticular dermis provide a route for nutrient supply and waste removal from the live cells in the VE. Traveling deeper into the dermis, hypodermis, and subcutaneous tissue, the blood vessels become larger and less dense, draining into the superficial venous plexus and the deep venous plexus. Stem and progenitor cells, located at the dermoepidermal junction, are responsible for skin renewal, continually pushing dead and dying cells into the SC and regenerating the underlying viable tissue.
The SC is the outermost layer, composed of corneocytes arranged in a “bricks and mortar” structure; it comprises the major physical barrier function of the skin (Figure 9.1). The corneocytes are terminally differentiated and flattened keratinocytes, rich in keratin filaments and embedded in a dense matrix of proteins to restrict transport of large molecules into the epithelial tissues beneath. The physicochemical properties of this layer are primarily responsible for limiting the free diffusion of molecules into the skin to those species <500 Da (approximately), which are moderately soluble in both water and oil phases.
The VE, i.e., the layers of the epidermis deep to the SC, is primarily composed of a densely packed layer of keratinocytes (≈95%) interspersed with small numbers of highly specialized cell types, including Langerhans cells (3%–5%; dendritic cells) and melanocytes (≈8%; melanin production to prevent UVB-associated damage) (Figure 9.1). Langerhans cells are potent antigen-presenting cells (APCs) and are thought to be key mediators of the improved immune responses observed in some transdermal vaccination techniques (see Section 9.3.3).
The dermis is a highly hydrated tissue composed largely of a collagenous extracellular matrix, containing a capillary network and a relatively lower density of cells, primarily fibroblasts and mast cells. Dermal dendritic cells (dDCs) have relatively recently been identified as a distinct APC population native to this layer, and their principal functions are still being determined, along with their interactions with other cells, tissue layers, and other components of the immune system. The vasculature of the dermis provides routes for systemic drug delivery, via circulating blood or lymphatic fluid.
FIGURE 9.1 Epidermal cell layers. (Designua/Shutterstock.com.)
9.1.2 ADVANTAGES OF THE TRANSDERMAL ROUTE FOR DRUG DEliVERY
In general terms, transdermal delivery methods are considered to be more acceptable to patients in comparison to more invasive delivery routes, including needle/syringe methods. This is because transdermal methods are relatively noninvasive, do not necessarily require trained health staff for administration, and “access” to viable delivery sites is rarely a concern due to the high number of available skin sites. Importantly, a number of transdermal techniques also allow the user or medical practitioner to control release rates into the tissue, from rapid delivery applications (e.g., local anesthetics) through to sustained-release periods (up to 1 week, e.g., pain medications).
For the delivery of drugs to local skin epithelial tissues, transdermal methods can clearly target the tissue of interest without off-target systemic effects. Transdermal delivery also has advantages for systemic delivery applications due to direct and/or indirect uptake of drugs from skin tissue fluid into the blood or lymphatic circulation, respectively. This potentially avoids first-pass liver metabolism, thereby requiring lower initial concentrations for therapeutic effects. Controlled-release strategies can also be used to control the drug delivery rate and pharmacodynamic profile. Furthermore, in the last decade or so, a clear narrative is emerging that transdermal methods provide unique access to a unique population of immunologically sensitive cells, which can be exploited for immunotherapeutic and vaccine applications.
9.1.3 LIMITING FACTORS FOR TRANSDERMAL DRUG DELIVERY
The key limitation in transdermal drug delivery, particularly for diffusion-based mechanisms, is that efficient delivery is limited to those compounds that possess the physicochemical properties that favor transepithelial transport. These properties are described in detail in Chapter 4 (Section 4.3.4); to summarize here for the transdermal route, this refers to low-molecular-weight (<500 Da approximately) compounds, with high partition coefficients and moderate solubility in both oil and water phases. This is due primarily to the physicochemical properties of the SC, which are crucial to maintain the skin’s barrier function. To overcome this limitation, a range of methods have been developed to enhance diffusion through the SC, or breach this outer layer completely, to gain direct access to the epithelial tissues underneath.
The other limitation, again particularly related to diffusion-based methods, is that attempts to increase delivery rates through the SC (direct breaching, enhanced permeability, etc.) are generally accompanied by local tissue irritation and/or inflammation. This has been problematic for the health-care providers, FDA, and patients, although it is becoming apparent that a minor degree of local irritation may enhance some therapeutic strategies, especially those that target the immune system.
9.2 MECHANICAL AND BIOLOGICAL PROPERTIES OF THE PRINCIPAL SKIN LAYERS
9.2.1 STRATUM CORNEUM: THE KEY MECHANICAL BARRIER
The SC is a semipermeable barrier that, owing to its variable mechanical properties, is challenging to breach in a minimally invasive manner, to target the underlying viable skin strata (whether it is the VE, dermis, or a combination of both). Mechanically, the SC is classified as a bioviscoelastic solid and shows highly variable properties. Obvious differences include the huge variation in thickness and composition with the skin site and the age of an individual. However, there are more subtle and equally important variations in SC properties to consider when configuring targeting methods.
For example, the SC mechanical breaking stress is strongly influenced by the ambient humidity/moisture content—the relative humidity range from 0% to 100% results in a decrease in excised human SC breaking stress from 22.5 to 3.2 MPa. Similarly, an increase in ambient temperature also results in an SC breaking stress decrease by an order of magnitude. More recently, with indentation studies using small probes (diameters of 2 and 5 μm) fitted to a nanoindenter (Kendall et al. 2007), we have found more complexity and variation in key SC, and underlying VE, mechanical properties. This work was then extended, first by moving the probes into the dermis (Crichton et al. 2011) and then also with more comprehensive measurements including a larger range of probe sizes (0.5–40 μm; Crichton et al. 2013).
Specifically, we found:
• As the probe radius is decreased, the elastic modulus increases (Figure 9.2a). Indeed, taking into consideration the work of others (e.g., Boyer et al. 2009), we found that decreasing the probe radius by three orders of magnitude (i.e., from 1000 to 1 μm, approximately), the corresponding elastic modulus also increased by three orders of magnitude (from 104 to 107 MPa).
• Significant differences in instantaneous elastic modulus between skin strata, also influenced by probe size (Figure 9.2b). In general, for a given probe size, the elastic modulus decreased from the SC to the VE, before increasing to the highest level in the dermis.
These and other sources of variability in the SC mechanical properties present challenges and fresh opportunities in configuring approaches to breach the SC in a minimally invasive manner and effectively deliver drugs/vaccines to the underlying layers. The other interesting point to note about the SC is that 0.1% of its area is comprised of appendages (hair follicle, sebaceous glands, and sweat ducts) that can provide “shunt” pathways for ions or polar molecules that cannot otherwise passively diffuse through the barrier. Indeed, since these units protrude into the dermis, they are served by capillary loops, which can serve to transport delivered cargoes into the systemic circulation. However, due to the very low surface area available at the skin surface, this route may only be applicable to drugs that require only minute quantities to generate the desired therapeutic response (e.g., gene therapies, vaccines).
FIGURE 9.2 Mechanical properties of the skin. (a) Elastic modulus of skin tissue measured by several groups; (b) elastic modulus of each discrete tissue layer using a small (1.9 μm radius, as shown) or large (6.6 μm radius) probe. (Adapted from Biomaterials, 34(8), Crichton, M.L., Chen, X.F., Huang, H. et al., Elastic modulus and viscoelastic properties of full thickness skin characterised at micro scales, 2087–2097. Copyright 2013, with permission from Elsevier; From Biomaterials, 32(20), Crichton, M.L., Donose, B.C., Chen, X. et al., The viscoelastic, hyperelastic and scale dependent behaviour of freshly excised individual skin layers, 4670–4681. Copyright 2011, with permission from Elsevier.)
9.2.2 VIABLE EPIDERMIS AND DERMIS
In the VE, the skin has evolved a highly competent immunological function, with an abundance of Langerhans cells (500–1000 cells mm−2), often serving as the first line of defense against many pathogens. As one example, Langerhans cells (Figure 9.1) are extremely effective APCs, responsible for the uptake and processing of foreign materials in order to generate an effective immune response. Such cells are reported to be up to 1000-fold more effective than keratinocytes, fibroblasts, and myoblasts at eliciting a variety of immune responses (Banchereau and Steinman 1998). Similarly, there are many types of APCs within the dermis that also serve important roles in immu-nological function.
The “target tissue” for most transdermal delivery approaches is the highly vascularized and immunologically rich dermal layer. In humans, the blood vessel density varies from 20 to 40 cm-2 (Huzaira et al. 2001), providing sufficient access to the systemic circulation for rapid uptake. Furthermore, the dermis contains a distinct population of immunologically active cell types in comparison to the VE. Less information is available on underlying dermal APCs. New populations of dDCs that express langerin (originally believed to be an exclusive marker for Langerhans cells) reportedly exist and have unique immunological functions within the skin. Thus, directly targeting specific Langerhans cells or dermal APC populations will allow immune responses to be modulated for immunotherapeutic purposes. One example is for potent immune responses for vaccines (Fernando et al. 2010; Sullivan et al. 2010), while allergy immunotherapy remains an area yet to be explored in detail.
9.3 APPROACHES TO TRANSDERMAL DRUG DELIVERY
The various approaches to achieving transdermal drug delivery can be grouped into three main areas (Figure 9.3):
1. Passive approaches, relying on diffusion
2. Active approaches, requiring the application of energy
3. Physical approaches, which breach the SC
Passive approaches rely on passive diffusion of the API through the skin layers, to reach the underlying capillaries of the dermis. As such, these approaches have been limited to relatively small-molecule and lipophilic active substances, since only these can penetrate the skin in therapeutically relevant amounts. However, significant research and innovation has resulted in the invention of a range of methodologies that overcome this limitation, and currently transdermal drug delivery methods exist that arguably cover the full range of human therapeutic molecules.
There is a general acceptance and approval of transdermal technologies for local drug delivery to the skin (e.g., anesthesia, hyperhidrosis). However, further studies are required for widespread approval of systemic delivery systems. The primary concern with the FDA appears to be related to areas that are yet to be thoroughly investigated—especially mechanistic information on skin irritation and minor burns associated with some techniques, and the lack of control over the delivered dose. With a needle/syringe system, the delivered dose is considered obvious (i.e., visual inspection of a metered syringe); however, the same cannot yet be stated for transdermal techniques. This is particularly important in terms of potential toxic effects related to overdose of some medications (e.g., fentanyl is a commonly studied drug for transdermal delivery systems in controlling pain medication, but as an opiate, it can also cause respiratory depression), based on patient-to-patient variability, and the differences between adults and children in pharmacokinetics and metabolism.
FIGURE 9.3 Summary of approaches to achieving transdermal drug/vaccine delivery. Passive approaches: (a) drug, (b) drug + penetration enhancer, (c) transdermal patch. Active approaches: (d) electroporation, (e) iontophoresis (arrow = induced convective flow, could be electroosmosis or electrophoresis), (f) ultrasonic wave. Physical approaches: (g) intradermal injection, (h) liquid jet injector, (i) gene gun. (j) microneedle arrays, delivering to either the dermis or the viable epidermis.
9.3.1 PASSIVE APPROACHES RELYING ON DIFFUSION
9.3.1.1 Transdermal Therapeutic Systems
Transdermal patches, also called transdermal therapeutic systems (TTS), comprise medicated skin patches that facilitate controlled release of the API (Figure 9.3c). They are designed so that the drug diffuses through the patch at a rate that is much slower than its diffusion-rate through the SC. Therefore drug is delivered at a controlled rate that is dictated by the patch, and not the skin.
The first TTS, developed by Alza (see Chapter 1, Section 1.4.1), was FDA-approved for use in humans in 1979 and comprised a skin patch that delivered scopolamine for up to 3 days (Transderm Scop®), for the treatment of nausea related to motion sickness. Nicotine patches followed approximately a decade later and remain the most well-known transdermal system for the general public. Structurally, a TTS contains a number of layers, including (1) an impermeable backing layer, (2) a layer (or layers) that provide the controlled-release mechanism, and (3) a peel-away strip, which is removed prior to application. Various mechanisms are used to provide controlled release, including (1) reservoir systems, whereby drug release is controlled by a rate-controlling membrane (RCM), which may be porous or microporous (Figure 9.4a), and (2) matrix systems, whereby drug release is controlled by diffusion through an adhesive matrix (Figure 9.4b).
Many other designs are used to facilitate controlled release, including, for example, a multilayered matrix TTS, which contains more than one drug-in-adhesive layer, the individual layers may be separated by a membrane. A reservoir-matrix TTS contains a drug reservoir layer with an associated RCM and an additional drug-in-adhesive matrix layer. More recently, TTS designs have moved away from reservoir systems to other CR mechanisms, thus avoiding adverse events related to rapid “dumping” of drug molecules into the skin, should the RCM become damaged.
The patches are recommended only for use for up to 7 days at the same location, to avoid local adverse skin responses mainly related to skin stripping upon removal (Wohlrab et al. 2011). All of the drugs approved for delivery using TTS technology (for example, nicotine, clonidine, scopolamine, nitroglycerin, and fentanyl) fit the required physicochemical profile described earlier, i.e., small and lipophilic molecules. The great challenge is that only a limited number of drugs are amenable to administration using this approach, so large and/or hydrophilic drugs are generally not deliverable using a TTS.
FIGURE 9.4 Transdermal therapeutic systems. Mechanisms of achieving controlled release: (a) reservoir system and (b) matrix system.
To improve the passive diffusion rates across the skin for a range of molecules, chemical enhancers have been introduced in order to increase skin permeability by disrupting the lipid structure of the SC (Figure 9.3b). However, increased permeation enhancement is also related to increased irritation and injury to the underlying layers. This approach provides enhanced permeability, and sometimes driving force, for transdermal transport. Hundreds of such enhancers have been trialed, with several specifically designed and approved for human use (1-dodecyl-azacycloheptan-2-one [Azone] and 2-n-nonyl-1,3dioxolane [SEPA™]). While a key challenge is identifying useful chemical enhancers for a wide range of different drugs, high-throughput screening approaches have been developed, using spectroscopic markers that correlate with (1) lipid permeability and (2) skin irritation. Amphiphiles containing either long-carbon tails or multiple aromatic rings were predicted to be efficient and indeed have been proven so experimentally, particularly by Mitragotri and colleagues.
A key limitation with conventional chemical enhancers is that their effects are often not limited to the SC layer itself; hence, irritation or injury to the underlying tissues often results. A range of alternative or combination approaches has been used with some success. Packaging of drugs inside micro- and nanosized carriers, such as liposomes, dendrimers, and nanoparticles, has the advantage of limiting delivery of the carrier to the SC alone so that enhancers do not reach the underlying tissue. Such encapsulation methods can have other advantages as well, including protective effects, additional control over release profiles, etc.
Biochemical enhancers have also been introduced, which comprise polypeptides that have pore-forming or pore-enhancing biological activity. These can often be formulated with conventional chemical enhancers, using the latter to increase lipid disruption at the SC.
Other combinations of enhancers have also been trialed in order to balance the permeability/irritation aspects of individual enhancers. In a high-throughput screening approach, one study examined 500 pairs of chemical enhancers in over 5000 individual compositions, finding a combination that provided significant permeability enhancement with low skin irritation (a combination of sodium laureth sulfate and phenylpiperazine at 0.35% and 0.15% by weight, respectively, in a 1:1 mixture of ethanol and phosphate buffer), validated in a rat model of leuprolide acetate peptide delivery (Karande et al. 2004).
Alternatively, some investigators have made use of the appendage delivery pathways to bypass the need to interact with the SC altogether—particularly for skin-specific and cosmetic treatments; however, this area is in its infancy (Chourasia and Jain 2009).
9.3.2 ACTIVE APPROACHES REQUIRING APPLICATION OF ENERGY
While conventional and novel chemical enhancers have resulted in significantly improved drug delivery in preclinical and clinical studies, resulting in FDA-approved therapies, this approach is still fundamentally limited to small molecule delivery. The application of electrical or ultrasonic energy to the skin has been shown to significantly increase the permeability of the SC above and beyond the activity of enhancers, often with minimal injury to underlying tissues. The high electrical and thermal resistance of the SC layer means that relatively high currents or temperature can be applied transiently, inducing significant permeability in the SC alone, with the effect lasting up to several hours following the initial energy input.
High voltage pulses (100–1500 V), applied transiently to the SC (0.01–10 ms), have been shown to disrupt lipid structure in the SC allowing drug transport via diffusion through “electropores” and electrophoresis-driven flow (Figure 9.3d). The level of enhancement in terms of delivered drug can be several orders of magnitude higher than control skin, depending on physicochemical properties of the drug of interest. However, this technique is still sensitive to the molecular weight of the drug to be delivered, its lipophilicity, competitive ions present in the formulation, and also the viscosity of the formulation. To date, most studies have been limited to preclinical animal studies due to the complexities involved in designing the devices (including proteins and DNA); however, both ascorbic acid and lidocaine have been successfully delivered to human volunteers via this method. Concern over safety and pain related to this technique are largely offset by the acceptance of related techniques (e.g., electrochemotherapy—8 × 1000 V cm-1, 100 μs with plate electrodes); however, further design improvements are required to limit electrical activity to the SC, thus avoiding excitation of sensory nerves in the deeper dermis (Denet et al. 2004).
Iontophoresis involves application of a low-voltage current in order to increase permeability but also provides an electrical driving force for transport (Figure 9.3e). An Ag/AgCl electrode set is favored due to the pH stability, as it does not produce the sharp decrease in pH that can be seen with Pt electrodes. The driving force is distinct to that induced by electroporation because it facilitates either direct transfer of charged moieties (electrophoresis) or mobile ion flows (electroosmosis), causing convective transport of weakly charged, or neutral, molecules. Interestingly, the delivery rate is a function of the current, therefore complex delivery profiles can be achieved for controlled-release purposes. Indeed, low-cost approaches to iontophoresis are available: simply linking the patch to a disposable, constant-voltage, disposable battery, avoiding the need for a microprocessor. For this reason, iontophoresis is considered suitable for clinical use because the process can easily be controlled by untrained individuals. While a range of devices have been commercialized and approved for local treatments (e.g., lidocaine for anesthesia, tap water for hyperhidrosis), there is still concern over systemic drug delivery (e.g., fentanyl for pain relief) due to potential side effects based on the variation in the delivered dose between patients, due to age or skin type (Kalia et al. 2004).
Ultrasound (also referred to as sonophoresis or phonophoresis) was first used as a skin permeabilization method for the local delivery of corticosteroids to the skin in the 1950s. The ultrasound wave is a compression wave (frequency >20 kHz), in which an electrical signal is transmitted to a “horn” and converted into a sound wave via piezoelectric crystals (change static dimensions in response to electric field) (Figure 9.3f). This causes the horn to be displaced in a cyclic fashion, with characteristic amplitude and frequency. Early studies utilized high-frequency sonophoresis (HFS; ≈0.7–16 MHz), resulting in up to 10-fold increase in local drug delivery through the SC. However, it was later recognized that low-frequency sonophoresis (LFS; ≈20–100 kHz) led to significant improvements in local delivery rates (up to three orders of magnitude better than HFS) due to the effects of generating, and collapsing, air bubbles near the SC (cavitation).
Key variables in ultrasound-induced drug delivery include the type of pulse program (e.g., continuous vs. timed pulses), the horn-to-skin distance, the overall treatment time, and the ultrasound medium composition. The medium is crucial: it contains the drug to be delivered, generally a chemical enhancer, and is usually formulated in aqueous solutions for LFS and gels for HFS, both of which are designed to have acoustic impedance similar to that of the skin. Pulsing is often used to reduce local heating of the tissue, although a small degree of heating can also act to help permeabilize the SC. The horn-to-skin distance is generally very close to (if not touching) the skin for HFS treatments, but often up to 1 cm from the skin for the LFS treatments. This is because the cavitation-associated effects of LFS treatment are dependent on the generation of gas bubbles in the medium above the skin, an effect diminished if the probe and skin are in direct contact. Treatment times can be wide ranging, from transient delivery regimes on the order of seconds/minutes up to steady-state delivery regimes, which can be on the order of hours/days. Synergistic effects have been observed when combining chemical enhancers into the medium (e.g., surfactants).
HFS treatments have found utility in the local delivery of low-molecular-weight (<1000 Da), skin-specific drugs, but generally are not favored for systemic delivery. In comparison, LFS treatments are still relatively new (only 20-year history of research so far); however, a number of groups have demonstrated that macromolecules, including proteins (e.g., interferon-γ ≈17 kDa, erythropoietin ≈48 kDa), can be effectively delivered via this method, paving the way for systemic delivery. Vaccines have also been delivered via this route (e.g., tetanus toxoid), with LFS treatment clearly showing activation of skin-resident APCs, paving the way for future innovations in transdermal vaccination. However, the prospect of systemic effects is also a safety issue; hence, the development pathway is expected to be longer for LFS treatments.
9.3.3 PHYSICAL APPROACHES THAT BREACH THE STRATUM CORNEUM
While the techniques described earlier have shown utility in permeabilizing the SC, allowing transport of mainly low-molecular-weight compounds via different mechanisms, the size and other physicochemical properties of the drug and/or formulation media are still key parameters that limit delivery and uptake by the key epidermal and dermal layers. To address this problem, physical methods of breaching the SC in order to deliver drug payloads directly in the epidermis and dermis have been developed. The key advantages of this approach include (1) the variety of drugs based on size/charge/polarity/hydrophilicity that can be delivered, due to the lack of direct interaction with the SC, and (2) in terms of vaccine delivery, compounds that can be delivered directly into the vicinity of the potent APCs present in the VE and dermis. For these reasons, physical approaches typically address the therapeutic challenge of vaccine delivery (with or without immune-boosting adjuvants) to the epidermal layers. Importantly, as the methods earlier have proven in many ways well suited to the local delivery of small, lipophilic compounds, physical approaches have historically been developed for a different challenge.
9.3.3.1 Intradermal Injections
Conceptually, the simplest method for targeting the epidermal layers of the skin with a drug is to simply inject the compound into that layer via a standard needle/syringe (Figure 9.3g). However, this approach proves to be extremely challenging from a technical point of view, due to the submillimeter precision required to target the appropriate layers. Furthermore, variations with individuals (injection site) and between individuals based on age, gender, levels of sun exposure, etc., can have dramatic effects on the hydration levels and hence mechanical properties of the skin layers, which cannot be predicted or controlled for, using such an approach. However, the intradermal injection route has certainly proven the concept that targeting immune cells in the skin results in significantly higher antibody titers for a range of vaccines (e.g., rabies, influenza), in comparison to the more traditional intramuscular or subcutaneous injection routes.
Interest in using high-speed liquid jet injectors arose in the mid-twentieth century because of its needle-free approach (Furth et al. 1995). Mitragotri provides an excellent review on the status of the liquid injector field (Mitragotri 2006). This technique has seen resurgence, with liquid delivered around the Langerhans cells in gene transfer and DNA vaccination experiments, and the delivery of drugs. Since then, there have been many developments in the field. This includes the licensing of new liquid jet injector devices for vaccine delivery and, separately, research investigations into microscale liquid jet injectors. As shown in Figure 9.3h, current liquid jet injectors typically disrupt the skin in the epidermal and dermal layer. To target exclusively the viable epidermal cells, such as Langerhans cells (as one example), the challenge of more controlled delivery needs to be addressed. With the dermal disruption induced by administration, liquid jet injectors are also reported to cause pain to patients.
With gene guns, pharmaceutical or immunomodulatory agents formulated as particles are accelerated in a supersonic gas jet to sufficient momentum to penetrate the skin (or mucosal) layer and to achieve a pharmacological effect (Figure 9.3i). Klein et al. pioneered this innovation with systems designed to deliver DNA-coated metal particles (of diameter of the order of 1 μm) into plant cells for genetic modification, using pistons accelerated along the barrels of adapted guns (Klein et al. 1987). The concept was extended to the treatment of humans, with particles accelerated by entrainment in a supersonic gas flow. Prototype devices embodying this concept have been shown to be effective, painless, and applicable to pharmaceutical therapies ranging from protein delivery to conventional vaccines and DNA vaccines.
Prior to operation, the gas canister is filled with helium or nitrogen to 2–6 MPa, and the vaccine cassette, comprising two 20 mm diaphragms, is loaded with a powdered pharmaceutical payload of 0.5–2 mg. The pharmaceutical material is placed on the lower diaphragm surface. Operation commences when the valve in the gas canister is opened to release gas into the rupture chamber, where the pressure builds up until the two diaphragms retaining the vaccine particles sequentially burst. The rupture of the downstream diaphragm initiates a shock that propagates down the converging–diverging nozzle. The ensuing expansion of stored gas results in a short-duration flow (often <1 ms) in which the drug particles are entrained and accelerated through the device. After leaving the device, particles impact on the skin and penetrate to the epidermis, to induce a pharmacological effect.
9.3.3.4 Microneedles, Microprojections, and Arrays Thereof
Microprojection arrays (MPAs), or “microneedles,” comprise arrays of sharp-tipped projects that are designed to pierce the tough SC layer and rapidly deliver vaccine payloads to the underlying epidermal and dermal tissues (Figure 9.3j). These devices emerged in the late 1990s and have now been fabricated from a range of materials (silicon, metal, polymer, ceramic, and hybrids/composites), in a range of different geometries (density of projections 101–104 cm-2; length 0.03–3 mm). Key design parameters include the shape, density, length, and tip sharpness of the projections, along with the application velocity, as these directly affect the skin penetration depth achieved by the array. The different materials of fabrication allow for different delivery modes, including (1) dissolution of a vaccine coating from a solid projection, (2) injection of a liquid vaccine through a hollow projection, and (3) projections designed to dissolve upon skin insertion, releasing the vaccine payload. Furthermore, coating strategies aim to optimize formulations for long-term thermostability (e.g., incorporating “glassy” sugars to stabilize biomolecules when dried), controlled release of vaccine payload following skin insertion, and high release efficiency into the skin (Prausnitz and Langer 2008; Kim et al. 2012).
Compelling data have been published showing that MPAs can elicit protective immune responses in comparison to standard intramuscular or subcutaneous injection in animal models, and in many cases with lower vaccine dosage required (e.g., Fernando et al. 2010). Influenza vaccine administration has been the most investigated test case using coated projections, showing protective immunity in comparison with the needle, on the basis of total IgG antibody levels and functional assays involving hemagglutination inhibition and neutralizing antibody activity (Kim et al. 2012). However, a range of different vaccines (including split virion, protein subunit, DNA plasmids, siRNA, etc.) targeting different diseases (influenza, human papillomavirus, West Nile virus, hepatitis B, herpes simplex virus, chikungunya virus, etc.) have been investigated.
To date, clinical trials have shown early promise both in terms of pain reduction and immunogenicity, in comparison to standard intramuscular injections for a range of approved and emerging vaccines. Most patients and clinicians report significantly lower pain scores for microneedle devices in comparison with standard hypodermic needles; there is minimal skin reactivity beyond a mild erythema, which resolves in minutes or hours in the majority of cases. Transdermal vaccination is discussed further in Chapter 17 (Section 17.4.2).
9.3.3.5 Nanopatch™: A Case Study
While early results in preclinical and clinical trials investigating microneedle technology are encouraging, key challenges remain to be solved, including minimizing delivered dose to increase distribution and availability and the significant improvement of vaccine formulation thermostability, for storage at ambient conditions for long periods. The Nanopatch™ (Vaxxas) was designed as an ultrahigh density MPA with vaccine formulated into a dry-coating layer over solid projections (Figure 9.5). The array design was fabricated in silicon, using standard semiconductor processing techniques. The coating rapidly dissolved once hydrated in the skin, for rapid delivery in seconds. In order to target the immune APCs in the mouse epidermis and dermis, a 21,000 cm-2 density array was designed, with very short tips (0.03–0.3 mm) in order to maximize the delivered dose at the target site. In a key proof-of-concept study, the authors showed that Nanopatches™ indeed targeted ≈50% of available APCs in mice (both in the VE and dermis), resulting in less than 100th of the standard intramuscular dose required for equivalent protective immune responses using an influenza-based mouse model (Fernando et al. 2010). Investigations have found that localized death of skin cells induced by the dynamic application of the Nanopatch plays an important role in generating potent immunity: a “physical immune enhancer” (Depelsenaire et al. 2014).
FIGURE 9.5 The Nanopatch™. (a–c) Arrays containing silicon projections of increasing length; (d–f) corresponding penetration of each array into mouse abdominal skin (cryo-SEM). (Adapted from Biomaterials, 34(37), Coffey, J., Corrie, S.R., and Kendall, M.A.F., Early circulating biomarker detection using a wearable microprojection array skin patch, 9572–9583. Copyright 2013, with permission from Elsevier.)
The Nanopatch™ has been used to formulate delivery strategies for a range of vaccines, and along the way, a number of novel coating technologies and formulations were developed to support long-term thermostability. Using a “jet-coating” approach, the Nanopatch™ technology has delivered many different classes of vaccine, including inactivated whole virus vaccines (e.g., FluVax®: commercially available seasonal influenza vaccine), viruslike particles (VLPs, e.g., Gardasil®: commercially available tetravalent HPV vaccine), and DNA plasmids (e.g., preclinical HSV2 vaccine), among others. Immune adjuvants, designed to boost the host immune response to a vaccine formulation, have also been incorporated into the coating mixtures to further enhance immune responses, e.g., from 100-fold to 900-fold dose reduction in a FluVax® model using Quil-A adjuvant (Fernando et al. 2010). Long-term thermostability has also been demonstrated using the jet-coating approach, with comparative immunogenicity observed with freshly coated devices or those coated and stored for over 6 months at 23°C prior to skin application.
In conclusion, transdermal drug delivery methods show great promise for both localized and systemic delivery of therapeutic molecules. Advances in technology over the past several decades have overcome the key problems encountered when attempting to deliver drugs through the SC, leading to the delivery of molecules with a wide variety of size, charge, and physicochemical properties. Most recently, technologies developed to breach the SC (while causing minimal discomfort) have allowed the delivery of vaccines and related macromolecules with/without adjuvants to the immunologically sensitive skin layers, leading to significantly improved vaccine responses in comparison to traditional methods (e.g., needle/syringe). These technologies (including gene guns, liquid jet injectors, and microneedles) will potentially allow more widespread distribution of life-saving vaccines and therapies to those who most need them, without the current problems arising from the need for needles, cold chains, and trained health-care professionals.
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