Historical Introduction to the Field of Controlled Drug Delivery


Historical Introduction to the Field of Controlled Drug Delivery

Anya M. Hillery and Allan S. Hoffman


1.1  Introduction

1.2  Early Drug Delivery Systems

1.3  The Spansule® Delivery System: The First Controlled-Release Formulation

1.4  Controlled Release Using a Rate-Controlling Membrane

1.4.1  Drug Diffusion through a Rate-Controlling Membrane

1.4.2  Osmotic Pressure Controlled Release: Water Diffusion through a Rate-Controlling Membrane

1.5  Long-Acting Injectables and Implants

1.6  Further Developments in Oral Controlled Release

1.7  Drugs on Surfaces

1.7.1  Heparinized Surfaces

1.7.2  Drug-Eluting Stents and Balloons

1.7.3  Mucoadhesives

1.8  Nanoscale DDS

1.8.1  Liposomes

1.8.2  Nanoparticle DDS

1.8.3  Polyplexes and Lipoplexes

1.8.4  Polymeric Micelles

1.9  Conclusions



This chapter presents a historical overview of the field of controlled drug delivery, describing how it grew in the past 60 years from a very small field, to the immense size and importance it represents today for human and animal health. This chapter also highlights many of the people who were involved in the conception and design of the key controlled drug delivery systems (DDS), as well as details about the compositions of the materials used. We begin by considering some of the earliest drug delivery formulations, followed by a discussion of some of the key technologies in the history of controlled drug delivery. It should be noted at the outset that in the early days of controlled drug delivery, the term “controlled release” tended to refer specifically to zero-order drug release obtained via a rate-controlling membrane (RCM), whereas the terms “sustained release” and “extended release” referred to the prolonged drug release obtainable using other DDS such as the oral Spansules® and bioerodible implants. With the passage of time, however, the delineation of these definitions has blurred. Currently, all these terms are used interchangeably, and the term “sustained release” is widely used.


Conventional oral delivery systems release the drug immediately in the lumen of the gastrointestinal (GI) tract. The drug then dissolves in the GI fluids and permeates the gut wall to be absorbed into the systemic circulation via the underlying blood capillaries. There is no control over the release of the drug.

An early example of modifying drug release via the oral route was the use of enteric coatings. Tablets can be coated with the so-called enteric polymers, which are nonswelling and hydrophobic at the acidity of the stomach, but become ionized, and then dissolve and release the drug, once they enter the slightly alkaline pH of the intestinal region of the GI tract. Thus, drug release is delayed from the stomach to the small intestine. These “delayed release” coatings are useful to either (1) protect the stomach from drugs that can cause gastric irritation (e.g., aspirin) or (2) protect drugs that can be destroyed in the acidic gastric environment (e.g., some penicillins). Early coatings, introduced in the late 1800s, such as keratin and shellac suffered from storage instability and also, crucially, the pH at which they dissolved was too high for adequate dissolution in the small intestine, so that they were not very effective.

In 1951, cellulose acetate phthalate was introduced as an enteric-coating material (Malm et al. 1951). This polymeric cellulose derivative dissolved at a very weakly alkaline pH, such as found in the small intestine, making it highly suitable for enteric controlled-release applications. Many enteric-coating products followed, including the commercially very successful poly(methacrylates), now marketed as the Eudragit® L and Eudragit® S series by Evonik Industries. Figure 1.1 shows compositions of some enteric-coating polymers.

With respect to parenteral delivery, the development of controlled-release systems began in the 1930s, with the introduction of compressed pellets of hydrophobic compounds, which could provide sustained drug release over time, thereby allowing a reduction in the dosing frequency. Pellets consisting of compressed, finely powdered, estradiol particles were administered via subcutaneous (s.c.) implantation to animals, to cause rapid weight gain in the treated animals. Subsequently, other pellet-type implants were developed using other steroidal hormones. The rate of sustained release of the hydrophobic drugs was determined by the relative hydrophobicity of the pellet (Chien 1982; Hoffman and Wright 2012).


FIGURE 1.1 Enteric-coating polymers.


Even though drug release could be delayed by using enteric coatings, these formulations still featured immediate release of the drug upon removal of the enteric coating. The next stage of technological development was the design of true controlled-release systems, designed to control the drug release rate throughout the lifetime of the formulation. The first of these was the Spansule oral DDS (Figure 1.2), introduced in 1952 by Smith, Kline & French (SKF) for the 12-hour delivery of dextroamphetamine sulfate (Dexedrine®). Each Spansule® capsule contains hundreds of tiny drug-loaded beads, coated with a variable layer of natural waxes, such as carnauba wax, beeswax, or glyceryl monostearate. On ingestion, the outer capsule rapidly disintegrates, liberating the drug-loaded beads. The waxy coating around the beads then gradually dissolves as they transit down the GI tract, to liberate the drug. The rate of drug release is controlled by the thickness and dissolution rate of the waxy coating. A single capsule contains subpopulations of beads with different coating thicknesses, to provide a sustained release of drug over time (Lee and Li 2010).

Subsequently, SKF introduced the cold remedy Contac® 600 (so called because each capsule contained 600 beads), which became the world’s leading cold or allergy remedy after its launch in 1960. Each capsule contained four distinct populations of beads: a quarter with no coating, for immediate drug release; a quarter with a thin waxy coating, which dissolved after 3 hours; a quarter had a thicker coating so that drug release occurred after 6 hours; and the final quarter, with the thickest coating, dissolved after 9 hours. In total, the beads from a capsule provided cold/allergy relief over a sustained 12-hour period. Many advertisements at the time described the Contac® system as “tiny little time pills,” which provided “12 hour cold or allergy relief,” thereby introducing the general public to the concept of sustained release (Figure 1.2). Since then, many drugs have been reformulated in the Spansule® system, although the original waxy coatings have largely been replaced with more stable and reproducible, slowly dissolving, synthetic polymers.


FIGURE 1.2 The Spansule system achieved “sustained” drug delivery kinetics over many hours.



Judah Folkman, an MD at Harvard University, was an early pioneer in the field of controlled drug delivery. He was circulating rabbit blood inside a Silastic® (silicone rubber [SR]) arteriovenous shunt, and when he exposed the tubing to hydrophobic anesthetic gases in the atmosphere surrounding the tubing, the rabbits went to sleep. He concluded that the gases were permeating across the SR tubing and absorbing into the blood. He proposed that sealed capsules of SR containing a drug could be implanted to act as a prolonged DDS (Folkman and Long 1964; Folkman et al. 1966;

Hoffman 2008).

In this way, a reservoir of drug is contained within a RCM. The drug can diffuse out through the reservoir at a controlled rate. If certain conditions are filled, drug release remains constant, “zero order” with time. The principle of the RCM zero-order DDS depends on a RCM that does not vary in permeation properties over the period of use. The zero-order condition also assumes that no significant diffusional resistances will be introduced with time, such as the deposition of a thick layer of scar tissue due to a foreign body response. Then, if the drug concentration–driving force from inside to outside of the device is constant, the delivery rate will be constant over the period of use (Figure 1.3).

Another key pioneer in the origin of the controlled drug delivery field was Alejandro Zaffaroni, a superb pharmaceutical chemist and entrepreneur who had collaborated with Carl Djerassi at Syntex on the synthesis of the steroid levonorgestrel, which was used in the first contraceptive pill. Zaffaroni had been thinking about creating a company devoted to controlled drug delivery. When he heard about Judah Folkman’s work, he went to visit him in Boston and Folkman agreed to become Chairman of the company’s Scientific Advisory Board. In 1968, Zaffaroni founded the very first company dedicated to the development of controlled drug delivery materials and devices, which he called Alza, after the first two letters of each of his first and last names. One of the authors (ASH) was invited to become a consultant at Alza and was thus a witness to the origins and growth of the controlled DDS field and met most of the pioneers personally over the years.


FIGURE 1.3 Membrane-controlled drug delivery systems. Zero-order delivery rate, controlled by a rate-controlling membrane.


FIGURE 1.4 Silicone rubber and poly(ethylene-co-vinyl acetate) were the first polymers to be used as rate-controlling membrane barriers, for the controlled delivery of small hydrophobic drugs.

The most common materials used as RCMs in the first devices were two polymers, SR and poly(ethylene-co-vinyl acetate) (EVA) (Figure 1.4). The EVA RCM is based on the copolymer of ethylene and vinyl acetate (VA). The VA disrupts the crystalline regions of the poly(ethylene) component, creating amorphous regions through which the drug can permeate (a drug cannot permeate through the crystalline region of a polymer). Thus, the higher the VA content of EVA, the higher the permeability of the drug through the EVA membrane. EVA RCMs may typically have as much as 40% VA.

A number of zero-order RCM DDS were developed in the 1970s and were approved for clinical use in the 1980s–1990s (Hoffman 2008). Typically, the drugs delivered were small and relatively hydrophobic, such as a variety of contraceptive steroids, as well as LHRH analogs (for treating prostate cancer) and pilocarpine (for treating glaucoma). Alza’s first commercial product, the eye insert, Ocusert®, received FDA approval in 1974. The device released the antiglaucoma drug, pilocarpine, at a constant rate in the eye for 1 week, using an EVA RCM (Figure 1.5a). An EVA RCM was also used in Alza’s intrauterine device (IUD), Progestasert®, approved in 1976, which provided zero-order controlled release of the contraceptive steroid progesterone, for over a month (Figure 1.5b).

In addition to the Alza Corp., others such as the Population Council, WHO, Upjohn/Pharmacia Pharmaceuticals, and Planned Parenthood were active in the development, approval process, and marketing of contraceptive drug devices. Norplant® is a controlled DDS birth control device that was developed by Sheldon Segal and Horatio Croxatto at the Population Council in New York in 1966 (Figure 1.5c). The original Norplant® consisted of a set of six small (2.4 mm × 34 mm) SR capsules, each filled with 36 mg of levonorgestrel (a progestin used in many birth control pills), for s.c. implantation in the upper arm. The implanted tubes had a 5-year duration of delivery, after which they had to be explanted surgically. Another set of six tubes could then be implanted if the patient desired it. Norplant® was approved for clinical trials in Chile in 1974 and finally approved from human use in Europe in the 1980s, followed by the United States in 1990. It was withdrawn by Wyeth Pharmaceuticals (who was marketing it in the United States) from the U.S. market in 2002 due to numerous “unwarranted” lawsuits. Production of the original Norplant® was discontinued globally in 2008.


FIGURE 1.5 Examples of early drug delivery devices based on a drug reservoir and a rate-controlling membrane, to effect controlled release. (a) Ocusert®, an eye insert, (b) Progestasert®, an intrauterine device, (c) Norplant®, for subcutaneous implantation, and (d) an early prototype of a vaginal ring. SR, silicone rubber; EVA, poly[ethylene-co-vinyl acetate].

In 2006, Organon introduced a single-tube system, Implanon®, using EVA as the RCM. The implant provides controlled release of the contraceptive drug etonogestrel for up to 3 years. More recently, Valera Pharmaceuticals introduced Vantas®, a tubular s.c. implant made of Hydron®, a poly(hydroxyethylmethacrylate) (polyHEMA) hydrogel. The implant provides continuous delivery, for over a year, of the gonadotropin-releasing hormone (GnRH) analog, histrelin acetate, for the treatment of prostate cancer. Similar to Norplant® and Implanon®, the Vantas® implant is nondegradable and has to be surgically retrieved after use.

Vaginal rings were also designed as zero order, RCM, DDS. Gordon Duncan at Upjohn/Pharmacia, with support of the WHO, developed an early example comprising a SR core, loaded with a contraceptive steroid and coated with SR (Figure 1.5d). Although it did not become commercialized, this ring laid the groundwork for the subsequent development of other vaginal rings, such as the Estring® and Femring®, which were approved in the late 1990s for the delivery of estradiol acetate, in the treatment of postmenopausal urogenital symptoms. NuvaRing®, developed at Merck, is made of EVA and has been used clinically to deliver estradiol for treating postmenopausal urogenital symptoms.

Zaffaroni was also interested in the potential of transdermal drug delivery. He patented the controlled delivery rate skin patch as a “Bandage for Administering Drugs” in 1971, shortly after he had founded Alza (Figure 1.6). The Alza skin patch is a reservoir system that incorporates two release mechanisms: an initial burst release of drug from the adhesive layer and zero-order release over an extended period (e.g., several days), facilitated by the RCM built into the patch and separating the drug reservoir from the skin surface. The skin patch technology is referred to as a transdermal therapeutic system (TTS). If the RCM does not change in properties during the contact time of the patch on the skin, the drug diffusion rate across the membrane and out of the patch will be constant. The delivery rate from the patch is designed to be much slower than the diffusion of the drug through the stratum corneum (the main resistance in the skin), thus rate control is determined by the patch and not the skin. This was referred to as “putting the major resistance to drug delivery into the device.”


FIGURE 1.6 Copy of the 1971 Alza patent for transdermal drug delivery using a skin patch (“bandage”).

Alza developed other types of RCMs that contain micropores, e.g., a stretched polypropylene (PP) membrane (Celgard®), with micropores that may be prefilled with mineral oil or wax. This membrane is used in some controlled-release skin patches. Since PP is highly crystalline, the wax- or oil-filled pores represent the preferred diffusion pathway for the drug. As long as the pore volume and pore interconnections within the microporous RCM do not change with the time of the patch on the skin, the drug diffusion rate across the membrane will be constant. As earlier, if the overall drug delivery rate from the patch is designed to be much slower than that by diffusion through the stratum corneum, the patch will exhibit a zero-order drug delivery rate.

The first controlled delivery skin patch commercially available was Alza’s product Transderm-Scop®, approved in 1979 for the transdermal delivery of scopolamine, a drug that alleviates the discomfort of motion sickness. It was developed with the idea that Alza could get funding from the U.S. space program for use of the patch in zero-gravity conditions. Many other TTS patches were subsequently developed by Alza and other companies, allowing once-a-day or even once-a-week dosing, with reduced side effects compared to the oral route (a further description of TTS is given in Chapter 9). It is important to note that some adhesives used to adhere the patch to the skin caused skin irritation.


In the 1970s, an alternative method to achieving controlled release was developed, based on the principles of the osmotic pump. It utilizes a constant volume and constant concentration (saturated) of a drug solution, or dispersion, inside a rigid, semipermeable membrane (the RCM). Water permeates through the RCM into the device, displacing an equal volume of drug solution out of the device, through a microscopic pore created in the membrane. The water permeates into the tablet due to an osmotic pressure difference between the osmotic pressure of water within the body fluids (e.g., the GI tract fluids) and the low osmotic pressure within the saturated drug condition inside. Figure 1.7 shows the elementary osmotic pump (EOP), developed by Felix Theeuwes and colleagues at Alza in 1975, for controlled-release oral drug delivery (Theeuwes 1975).


FIGURE 1.7 The elementary osmotic pump. Δπ = the osmotic gradient, i.e., the difference between the osmotic pressure in the surrounding environment (Δπe) and the osmotic pressure inside the device (Δπi).

It is important to emphasize that, while these devices exhibit zero-order drug delivery, they operate on a completely different delivery mechanism from the diffusion-driven, RCM devices described earlier. For osmotic pressure control, the constant drug delivery rate is driven by a membrane-controlled, constant rate of water permeation into the device (in contrast to drug diffusion out

May 8, 2017 | Posted by in PHARMACY | Comments Off on Historical Introduction to the Field of Controlled Drug Delivery

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