Rate Control in Drug Delivery and Targeting: Fundamentals and Applications to Implantable Systems

Rate Control in Drug Delivery and Targeting:

Fundamentals and Applications to Implantable Systems

Hongkee Sah and Yie W.Chien

4.1	Introduction	73
4.2	Advantages and disadvantages of implantation therapy	74
4.3	Biocompatibility issues	76
4.4	Non-degradable polymeric implants	77
4.5	Biodegradable polymeric implants	88
4.6	Implantable pumps	96
4.7	Conclusions	102
4.8	Further reading	102
4.9	Self-assessment questions	103

OBJECTIVES

On completion of this chapter the reader should be able to:

• Understand the advantages and disadvantages of implant therapy

• Describe the different types of non-degradable polymeric implants

• Describe the different types of biodegradable polymeric implants

• Describe rate control in drug delivery and targeting

• Give some examples of implant systems presently used in drug delivery

• Give examples of osmotic implant systems

4.1 INTRODUCTION

mplan adalah sistem penghantaran obat satu unit yang telah dirancang untuk mengirimkan sebagian obat dengan kecepatan yang diinginkan secara terapeutik, selama periode waktu yang lama. Sistem seperti itu paling sering digunakan untuk administrasi parenteral berkelanjutan, termasuk pemberian obat mata dan subkutan. Bab ini berfokus pada sistem implan dan mekanisme kontrol laju yang membentuk komponen intrinsik dari sistem implan. Karena mekanisme kontrol kecepatan ini dapat diterapkan pada banyak sistem penghantaran obat lain, bab ini juga berfungsi sebagai pengantar umum untuk metode kontrol kecepatan yang dapat dicapai dengan menggunakan pengiriman obat yang canggih dan strategi penargetan

74.

Implants are available in many forms, including:

• polymers, which can be biodegradable or non-degradable and are available in various shapes (rod, cylinder, ring, film, etc.), sizes and mechanisms of drug release;

• mini-pumps, which can be powered by osmotic or mechanical mechanisms.

Implan membutuhkan administrasi khusus untuk memulai terapi. Mereka biasanya ditanamkan secara subkutan, baik ke dalam jaringan interstisial lepas dari permukaan luar lengan atas, permukaan anterior paha atau bagian bawah perut. Namun, implan juga dapat dipasang melalui pembedahan, misalnya, rongga vitreous mata (implan intravitreal), atau intraperitoneal.

4.1.1 Historical development of implants

ada akhir tahun 1930-an, pelet yang terdiri dari partikel estradiol bubuk halus yang dikompresi ditanamkan secara subkutan pada hewan, yang menyebabkan hewan bertambah berat pada kecepatan yang jauh lebih cepat daripada hewan tanpa implan. Para ilmuwan selanjutnya membuat implan tipe pelet yang terdiri dari hormon steroid lain termasuk testosteron, progesteron, deoksikortikosteron dan dromostanolon propionat.

Pelepasan dari implan jenis pelet tersebut diatur oleh pelarutan sebagian obat tertentu dalam cairan tubuh dan dengan demikian tidak dapat diterima oleh kontrol eksternal. Implan tipe pelet juga tidak memiliki kemampuan reproduksi pelet-ke-pelet dalam kecepatan pelepasan obat. Oleh karena itu, upaya dilakukan untuk mengoptimalkan pendekatan tersebut. Pada awal 1960-an, dilaporkan bahwa senyawa dengan berat molekul kecil hidrofobik meresap melalui kapsul karet silikon dengan kecepatan yang relatif rendah. Ketika ditanamkan pada hewan, sistem melepaskan obat dengan kecepatan yang cukup konstan dan juga menimbulkan sedikit peradangan di tempat implantasi. Penggunaan elastomer silikon sebagai penghalang difusi untuk mengontrol pelepasan senyawa seperti hormon steroid, insektisida, anestesi dan antibiotik kemudian didemonstrasikan. Tingkat pelepasan obat tunduk pada kontrol eksternal dengan memanipulasi ketebalan, luas permukaan, geometri dan komposisi kimia dari elastomer silikon.

Karena membran karet silikon tidak permeabel terhadap senyawa hidrofilik atau dengan berat molekul tinggi, upaya bersama dilakukan untuk mengembangkan polimer biokompatibel lainnya untuk digunakan dalam perangkat implan. Polimer tersebut termasuk poli (etilen-ko-vinil asetat), poli (etilen), poli (propilena), poli (hidroksimetil metakrilat), poli (laktida-ko-glikolida), poli (anhidrida) dan poli (orto ester). Karakteristik dan aplikasi masing-masing keluarga polimer penting akan dibahas nanti di bab ini

4.2 ADVANTAGES AND DISADVANTAGES OF IMPLANTATION THERAPY

Implants possess several advantages, but also disadvantages, as drug delivery systems depending on the nature of the drug being delivered. A brief overview of both the advantages and disadvantages of implantable drug delivery is given below.

4.2.1 Advantages

The advantages of implantation therapy include:

• Kenyamanan: konsentrasi obat yang efektif dalam aliran darah dapat dipertahankan untuk waktu yang lama metode seperti infus intravena terus menerus atau suntikan yang sering. Namun, di bawah rejimen ini, pasien seringkali diharuskan untuk tinggal di rumah sakit selama pemberian untuk pemantauan medis berkelanjutan. Obat kerja pendek memperburuk situasi, karena jumlah suntikan atau kecepatan infus harus ditingkatkan, untuk mempertahankan tingkat obat yang efektif secara terapeutik. Sebaliknya, terapi implantasi memungkinkan pasien untuk menerima pengobatan di luar lingkungan rumah sakit, dengan pengawasan medis yang minimal. Terapi implantasi juga ditandai dengan insiden komplikasi terkait infeksi yang lebih rendah dibandingkan dengan sistem infus berbasis kateter.

• Kepatuhan: dengan memungkinkan pengurangan, atau penghapusan total, dosis yang melibatkan pasien, kepatuhan meningkat sangat. Seseorang bisa saja lupa minum tablet, tetapi pemberian obat dari implan sebagian besar tidak tergantung pada masukan pasien. Beberapa sistem implan melibatkan pengisian ulang berkala, tetapi meskipun faktor ini pasien kurang terlibat dalam memberikan obat yang diperlukan.

• Potensi untuk pelepasan terkontrol: tersedia implan yang mengirimkan obat-obatan dengan perintah nol terkontrol melepaskan kinetika. Sebagaimana dibahas dalam Bab 1 (Bagian 1.5.1), rilis terkontrol orde-nol menawarkan keuntungan dari:

(i) avoiding the peaks (risk of toxicity) and troughs (risk of ineffectiveness) of conventional therapy;

(ii) reducing the dosing frequency;

(iii) increasing patient compliance.

• Potensi pelepasan intermiten: pompa yang dapat diprogram secara eksternal (dibahas nanti dalam bab ini) bisa memfasilitasi rilis intermiten. Sebagaimana dibahas dalam Bab 1 (Bagian 1.5.2), pelepasan intermiten dapat memfasilitasi pelepasan obat sebagai respons terhadap faktor-faktor seperti:

(iv) circadian rhythms;

(v) fluctuating metabolic needs;

(vi) the pulsatile release of many peptides and proteins.

• Potensi pelepasan bio-responsif / pelepasan:bio-responsif dari implan merupakan area penelitian yang sedang berlangsung dan dibahas di Bab 16.

• Peningkatan pengiriman obat: menggunakan sistem implan, obat dikirim secara lokal atau sistemik sirkulasi dengan gangguan minimal oleh hambatan biologis atau metabolik. Misalnya, bagian obat melewati saluran pencernaan dan hati. Efek bypass ini terutama bermanfaat untuk obat yang diserap dengan buruk atau mudah dinonaktifkan di saluran gastrointestinal dan / atau hati sebelum distribusi sistemik.

• Fleksibilitas: fleksibilitas yang cukup dimungkinkan dengan sistem ini, dalam pemilihan bahan, metode manufaktur, tingkat pemuatan obat, laju pelepasan obat, dll.

• Komersial: bentuk sediaan implan mendiversifikasi portofolio produk obat tertentu (lihat Bagian 2.2). Dari perspektif regulasi, ini dianggap sebagai produk obat baru dan dapat memperluas perlindungan pasar obat selama 5 tahun tambahan (untuk badan obat baru) atau 3 tahun (untuk obat yang sudah ada)..

4.2.2 Disadvantages

The disadvantages of implantation therapy include such factors as:

• Invasive: as described in Section 3.5.2, either a minor or a major surgical procedure is required to initiate therapy. This requires the appropriate surgical personnel, and may be traumatic, time-consuming, cause some scar formation at the site of implantation and, in a very small portion of patients, may result in surgery-related complications. The patient may also feel uncomfortable wearing the device.

• Termination: non-biodegradable polymeric implants and osmotic pumps must also be surgically retrieved at the end of treatment. Although a biodegradable polymeric implant does not require surgical retrieval, its continuing biodegradation makes it difficult to terminate drug delivery, or to maintain the correct dose at the end of its lifetime.

• Danger of device failure: there is a concomitant danger with this therapy that the device may for some reason fail to operate, which again requires surgical intervention to correct.

• Limited to potent drugs: the size of an implant is usually small, in order to minimize patients’ discomfort. Therefore, most systems have a limited loading capacity, so that often only quite potent drugs, such as hormones, may be suitable for delivery by implantable devices.

• Possibility of adverse reactions: the site of implantation receives a high concentration of the drug delivered by an implant. This local high drug concentration may trigger adverse reactions.

• Biocompatibility issues: concerns over body responses to a foreign material often raise the issues of biocompatibility and safety of an implant (discussed in the next section).

• Commercial disadvantages: developing an implantable drug delivery system requires an enormous amount of R&D investment in terms of cost, effort, and time. If a new biomaterial is proposed to fabricate an implant, its safety and biocompatibility must be thoroughly evaluated to secure the approval of regulatory authorities. These issues can attribute to significant delay in the development, marketing and cost of a new implant.

4.3

BIOCOMPATIBILITY ISSUES

Implants may cause short- and long-term toxicity, as well as acute and chronic inflammatory responses.

Adverse effects may be caused by:

• The intact polymer: this may be due to the chemical reactivity of end or side groups in a polymer, organometallics used as polymerization initiators, or extractable polymeric fragments.

• Residual contaminants: such as residual organic solvents, unreacted monomers and additives used as fillers.

• Toxic degradation products: this effect is applicable to biodegradable polymers; for example, degradation of poly(alkylcyanoacrylate) leads to the formation of formaldehyde which is considered toxic in humans. In the case of a bioerodible poly(vinylpyrrolidone), the accumulation of the dissolved polymer in the liver raises a longterm toxicity issue.

• Polymer/tissue interfacial properties: the implant interface is a unique site where different chemicals co-exist and interact. If the surface of an implant has an affinity towards specific chemicals, an abnormal boundary layer will develop. The subsequent intra-layer rearrangement or reactions with other species then trigger tissue reactions. The defence reactions of the host tissue often lead to encapsulation of an

implant by layers of fibrous tissues. Since the encapsulation frequently impedes drug release, in vitro drug release data may not permit the prediction of in vivo drug release patterns. High local drug concentrations at the site of implantation over extended periods of time can also cause severe local irritation or adverse tissue reactions.

The performance and response of the host toward an implanted material is indicated in terms of biocompatibility. Major initial evaluation tests used to assess the biocompatibility of an implant are listed in Table 4.1. These tests include:

• observation of the implant/tissue interactions at the site of implantation;

Table 4.1 Examples of major initial tests for assessing the biocompatibility of an implant

Biological Effect	Prolonged Contact^a		Permanent Contact^b
	Tissue/Bone^c	Blood	Tissue/Bone	Blood
Cytotoxicity	x	x	x	x
Sensitization	x	x	x	x
Irritation or	∆	x	∆	x
intracutaneous
reactivity
Systemic toxicity	∆	x	∆	x
(acute toxicity)
Subchronic	∆	x	∆	x
toxicity (subacute
toxicity)
Chronic toxicity			x	x
Genotoxicity	x	x	∆	x
Implantation	x	x	x	x
Haemocompatibil		x	x	x
ity
Carcinogenicity			x	x

Source: FDA General Program Memorandum #G95-1 ^aContact duration ranges from 24 hours to 30 days.

^bContact duration is longer than 30 days. Tissue includes tissue fluids and subcutaneous spaces, X:

ISO (International Standards Organizations) evaluation tests for consideration. ∆: Additional tests

which may be applicable

• assessment of the intensity and duration of each inflammatory response;

• histopathological evaluation of the tissues adjacent to the implant.

4.4

NON-DEGRADABLE POLYMERIC IMPLANTS

Non-degradable polymeric implants are divided into two main types (see also section 3.2):

• reservoir devices, in which the drug is surrounded by a rate-controlling polymer membrane (which can be non-porous, or microporous);

• matrix devices, in which the drug is distributed throughout the polymer matrix.

In both cases, drug release is governed by diffusion, i.e. the drug moiety must diffuse through the polymer membrane (for a reservoir device) or the polymeric matrix (for a matrix device), in order to be released.

The choice of whether to select a reservoir-type, or a matrix-type, implantable system depends on a number of factors, including:

• the drug’s physicochemical properties;

• the desired drug release rate;

• desired delivery duration;

• availability of a manufacturing facility.

For example, it is generally easier to fabricate a matrix-type implant than a reservoir system, so this may determine the selection of a matrix system. However, if drug release is the overriding concern, a reservoir system may be chosen in preference to a matrix system. This is because reservoir systems can provide zero-order controlled release, whereas drug release generally decreases with time if a matrix system is used.

4.4.1

Reservoir-type non-degradable polymeric implants

4.4.1.1

Solution diffusion

For solution diffusion, a drug reservoir is bound by a polymeric membrane which has a compact, non-porous structure and functions as a rate-controlling barrier (Figure 4.1).

Silicones are used extensively as nondegradable non-porous membranes. They are polymerized from siloxanes and have repeating OSi(R₁R₂) units. They vary in molecular weight, filler content, R₁ and R₂, and the type of reactive silicone ligands for cross-linking. Variations in these parameters permit the synthesis of a wide range of material types such as fluids, foams, soft and solid elastomers (Figure 4.2).

Poly(ethylene-co-vinyl acetate) (EVA copolymer) is also widely used as a non-degradable polymeric implant. These copolymers have the advantages of:

• Ease of fabrication: the copolymers are thermoplastic in nature, thus an implantable device is easily fabricated by extrusion, film casting or injection molding.

• Versatility: the copolymers are available in a wide range of molecular weights and ethylene/vinyl acetate ratios. As the ethylene domain is crystalline, an increase in the content of ethylene unit affects the crystallinity and the solubility parameter of the copolymer. Thus the release rate of a drug from the device can be tailored as required.

Other polymeric materials commonly used as non-porous, rate-controlling membranes are given in Table 4.2.

The penetration of a solvent, usually water, into a polymeric implant initiates drug release via a diffusion process. Diffusion of drug molecules through non-porous polymer membranes depends on the size of the drug molecules and the spaces available between the polymeric chains. Even through the space between the polymer chains may be smaller than the size of the drug molecules, drug can still diffuse through the polymer chains due to the continuous movement of polymer chains by Brownian motion.

For transport through the membrane, there are three barriers to be circumvented (Figure 4.3):

Figure 4.1 Reservoir-type polymeric implant

Figure 4.2 Structure of silicones (a) silicone fluid (Dow Corning 360 Medical Fluid); (b) silicone foam elastomer; (c) silicone elastomer (vulcanized Silastic 382 Medicalgrade Elastomer); and (d) silicone elastomer (vulcanized Silastic Medical Adhesive Type A)

Table 4.2 Polymers used for fabrication of reservoir systems

Polymers providing solution-diffusion mechanism

Silicone rubber, especially polydimethyl siloxane (Silastic)

Silicone-carbonate copolymers, Surface-treated silicone rubbers

Poly (ethylene-vinyl acetate), Polyethylene, Polyurethane (Walopur)

Figure 4.3 The steady-state concentration profile of a drug in a reservoir-type polymeric implant C_r=concentration of drug in the reservoir, C_i=concentration of drug at the site of implantation

Polymers providing solution-diffusion mechanism

Polyisopropene, Polyisobutylene, Polybutadiene

Polyamide, Polyvinyl chloride, Plasticized soft nylon

Highly cross-linked hydrogels of polyhydroxyethyl methacrylate,

Polyethylene oxide, Polyvinyl alcohol, or Polyinyl pyrrolidone

Cellulose esters, Cellulose triacetate, Cellulose nitrate

Modified insoluble collagen

Polycarbonates, Polyamides, Polysulfonates

Polychloroethers, Acetal polymers, Halogenated polyvinylidene fluoride

Loosely cross-linked hydrogels of polyhydroxyethyl methacrylate,

Polyethylene oxide, Polyvinyl alcohol or Polyvinyl pyrrolidone

• the reservoir-membrane interface;

• the rate-controlling membrane;

• the membrane-implantation site interface.

The drug molecules in the reservoir compartment initially partition into the membrane, then diffuse through it, and finally partition into the implantation site. The rate of drug diffusion follows Fick’s Law (see Section 1.3.3.2):

(Equation 4.1)

where dm/dt=the rate of drug diffusion

D=the diffusion coefficient of the drug in the membrane k=the partition coefficient of the drug into the membrane

h=the membrane thickness

A=the available surface area

∆C=the concentration gradient, i.e. C_r−C_i where Cr and C_i denote the drug concentrations in the reservoir and at the site of implantation respectively.

As sink conditions apply;

hence

(Equation 4.2)

Substituting further:

(Equation 4.3)

where P, the permeability constant, is defined as Dk/h and has the units cm/s. The release rate of a drug from different polymeric membranes can be compared from the corresponding P values.

Substituting again:

(Equation 4.4)

where K₁ is a pseudo-rate constant and is dependent on the factors D, A, k and h. This is the familiar form of a first-order rate equation and indicates that the rate of diffusion is proportional to drug concentration.

However, in this system, the drug reservoir consists of either:

• solid drug particles, or

• a suspension of solid drug particles in a dispersion medium

so that the concentration of drug (C_r) in the system always remains constant, so that Equation 4.4 simplifies to:

(Equation 4.5)

where K₂ is a constant and is dependent on C_r.

Equation 4.5 is the familiar form of a zero-order rate equation and indicates that the drug release rate does not vary with time (Figure 4.4). Thus the release rate of a drug from this type of implantable device is constant during the entire time that the implant remains in the body.

4.4.1.2

Pore-diffusion

In some cases, the rate-controlling polymeric membrane is not compact but porous. Microporous membranes can be prepared by making hydrophobic polymer membranes in the presence of water-soluble materials such as poly(ethylene glycol), which can be subsequently removed from the polymer matrix by dissolving in aqueous solution. Cellulose esters, loosely cross-linked hydrogels and other polymers given in Table 4.2 also give rise to porous membranes.

In microporous reservoir systems, drug molecules are released by diffusion through the micropores, which are usually filled with either water or oil (e.g. silicone, castor and olive oil). Solvent-loading of a porous membrane device is achieved simply by immersing the device in the solvent. When this technique presents some difficulty, the implantable device is placed inside a pressure vessel and pressure is then applied to facilitate the filling of the solvent into pores. The transport of drug molecules across such porous

Figure 4.4 “Mt” Zero-order controlled release profile of a reservoir-type nondegradable polymeric implant (porous or compact membrane)

membranes is termed pore-diffusion. The selection of a solvent is obviously of paramount importance, since it affects drug permeability and solubility.

In this system, the pathway of drug transport is no longer straight, but tortuous. The porosity ε of the membrane and the tortuosity τ of the pathway must therefore also be considered. Thus for a porous polymeric membrane, Equation 4.4 is modified as follows:

(Equation 4.6)

where C_s, the drug solubility in a solvent, is the product of K and C_r and D_s is the drug diffusion coefficient in the solvent.

As for the non-porous reservoir device, in the microporous system, both:

• the surface area of the membrane and

• the drug concentration in the reservoir compartment remain unchanged, thus “Mt” kinetics is again demonstrated and zero-order controlled release is attained (Figure 4.4).

4.4.1.3

Examples of non-degradable reservoir devices

Norplant subdermal implant

The Norplant contraceptive implant is a set of six flexible, closed capsules made of a dimethylsiloxane/ methylvinylsiloxane copolymer containing levonorgestrel. The silicone rubber copolymer serves as rate-

Figure 4.5 Structure of Vitrasert implant

controlling membrane. The capsules are surgically implanted subdermally, in a fan-like pattern, in the mid-portion of the upper arm. The implant releases levonorgestrel continuously at the rate of 30 µg/day (the same daily dose provided by the oral uptake of the progestin-only minipill) over a 5-year period. After the capsules are removed, patients are promptly returned to normal fertility.

Vitrasert intravitreal implant

The Vitrasert implant has been developed to deliver therapeutic levels of ganciclovir locally to the eye, for the treatment of retinitis infected by Cytomegalovirus (CMV) (see Section 12.4.2). Localized delivery to the eye minimizes the systemic side effects of the drug. The implant is surgically placed in the vitreous cavity of the eye and delivers therapeutic levels of ganciclovir for up to 32 weeks.

The implant consists of a tablet-shaped ganciclovir reservoir. The drug is initially completely coated with poly(vinyl alcohol) (PVA) and then coated with a discontinuous film of hydrophobic, dense poly (ethylene-co-vinyl acetate) (EVA). Both polymers are nonerodible and hydrophobic (the PVA used in the implant is cross-linked and/or high molecular weight, to ensure it does not dissolve when exposed to water). The entire assembly is coated again with PVA to which a suture tab made of PVA is attached (Figure 4.5).

The first step for drug release involves the dissolution of ganciclovir by ocular fluids permeating through the PVA and EVA membranes. The drug molecules permeate through the PVA membrane, then through the pores of the discontinuous film of EVA and finally through the outer PVA membrane into the vitreous cavity, at the rate of approximately 1 µg/hr over a 7- to 8-month period. The release rate can be further tailored by varying the membrane characteristics of PVA and EVA.

4.4.2

Matrix-type non-degradable polymeric implants

In a matrix-type implant the drug is distributed throughout a polymeric matrix (Figure 4.6).

Matrix-type implants are fabricated by physically mixing the drug with a polymer powder and shaping the mixture into various geometries (e.g. rod, cylinder, or film) by solvent casting, compression/injection molding or screw extrusion.

The total payload of a drug determines the drug’s physical state in a polymer:

• Dissolved: the drug is soluble in the polymer matrix. A dissolved matrix device (also known as a monolithic solution) appears at a low payload.

• Dispersed: the drug is present above the saturation level, additional drug exists as dispersed particles in the polymer matrix (also known as a monolithic dispersion).

Figure 4.6 Matrix-type polymeric implant

• Porous: with further increase in total drug payload, the undissolved drug particles keep in contact with one another. When the drug content occupies more than 30% volume of the polymer matrix, the leaching of drug particles results in the formation of pores or microchannels that are interconnected.

Regardless of a drug’s physical state in the polymeric matrix, the release rate of the drug decreases over time. Initially, drug molecules closest to the surface are released from the implant. As release continues, molecules must travel a greater distance to reach the exterior of the implant and thus increase the time required for release (Figure 4.7). This increased diffusion time results in a decrease in the release rate from the device with time (Figure 4.8). Numerous equations have been developed to describe drug release kinetics obtainable with dissolved, dispersed, and porous-type matrix implants, in different shapes, including spheres, slabs and cylinders. Suffice to say here that in all cases, the release rate initially decreases proportionally to the square root of time:

(Equation 4.7)

where k_d is a proportionality constant dependent on the properties of the implant, thus:

(Equation 4.8)

This “Mt^1/2” release kinetics is observed for the release of up to 50–60% of the total drug content.

Thereafter, the release rate usually declines exponentially.

Thus a reservoir system can provide constant release with time (zero-order release kinetics) whereas a matrix system provides decreasing release with time (square root of time-release kinetics). A summary of the drug release properties of reservoir and matrix nondegradable devices in given in Table 4.3.

The decreasing drug release rate with time of a matrix system can be partially offset either by:

• designing a special geometry that provides increasing surface over time (this strategy is used in the Compudose implant, described in Section 4.4.2.1 below), or

• using reservoir/matrix hybrid-type systems (this strategy is used in the Synchro-Mate-C and Implanon implants, described in Section 4.4.3).

Table 4.3 A summary of the drug release properties of reservoir and matrix nondegradable implant devices

System	Release Mechanism	Release Properties	Release Kinetics

Reservoir	Diffusion through a polymeric	Constant drug release with time	Zero-order release “M t”
	membrane (which can be
	compact or microporous)

Figure 4.7 A matrix-type implant in which a drug is dissolved. The initial diffusion of drug molecules leaves a drug-depleted polymeric zone with a length h, which increases with time. This event leads to an increase in diffusional distance over time

System	Release Mechanism	Release Properties	Release Kinetics

Matrix	Diffusion through a polymeric	Drug release decreases with time	Square root of time release “M t^1/
	matrix		2_”

4.4.2.1

Examples of matrix-type implants

Compudose cattle growth implant

In the Compudose implant microcrystalline estradiol is dispersed in a silicone rubber matrix, which is then used to coat a biocompatible inert core of silicone rubber, that does not contain any drug particles (Figure 4.9). This particular design, consisting of a thin layer of the drug-containing matrix and a relatively thick drug-free inert core, minimizes tailing in the drug release profile.

When this implant is placed under the skin of an animal, estradiol is released and enters into systemic circulation. This stimulates the animal’s pituitary gland to produce more growth hormone and causes the animal to gain weight at a greater rate. At the end of the growing period, the implant can be easily removed to allow a withdrawal period before slaughter.

The Compudose implant is available with a thick silicone rubber coating (Compudose-400) and releases estradiol over 400 days, whereas one with a thinner coating (Compudose-200) releases the drug for up to 200 days.

Figure 4.8 Drug release by diffusion through a nondegradable polymeric matrix. There is a decrease in the release rate from the device with time

Figure 4.9 Structure of Compudose cattle growth implant

Syncro-Mate-B implant

The implant consists of a water-swellable Hydron (cross-linked ethylene glycomethacrylate) polymer matrix in which estradiol valerate (Norgestomet) crystals are dispersed. It is used for the synchronization of estrus/ovulation in cycling heifers. Once implanted in the animal’s ear, the implant delivers estradiol valerate at the rate of 504 µg cm⁻² day^−1/2 over a period of 16 days.

Figure 4.10 Hybrid-type polymeric implants (a) Syncro-Mate-C: matrix containing microreservoirs of drug, (b)

Implanon: membrane coating a drug containing matrix

4.4.3

Reservoir/matrix hybrid-type polymeric implants

Reservoir/matrix hybrid-type non-degradable polymeric implants are also available. Such systems are designed in an attempt to improve the “Mt^1/2” release kinetics of a matrix system, so that release approximates the zero-order release rate of a reservoir device. Examples of these types of systems include:

Syncro-Mate-C subdermal implant

To make this implant, an aqueous solution of PEG is first loaded with estradiol valerate (Norgestomet) at a saturation level. This suspension is then dispersed in a silicone elastomer by vigorous stirring. The mixture is blended with a cross-linking agent, which results in the formation of millions of individually sealed microreservoirs. The mixture is then placed in a silicone polymer tube for in situ polymerization and molding. The tube is then sectioned to make tiny cylindrical implants (Figure 4.10a). Drug molecules initially diffuse through the microreservoir membrane and then through the silicone polymer coating membrane. This implant provides zero-order release kinetics, rather than square root of time-release kinetics. The two open ends of the implant do not affect the observed zero-order release pattern because their surface area is insignificant compared to the implant’s total surface area.

Implanon (Organon)

Implanon is fabricated by dispersing the drug, 3-ketodesogestrel, in an EVA copolymer matrix. This polymer matrix is then coated with another EVA copolymer, which serves as a rate-controlling membrane (Figure 4.10b). The drug permeation through the polymer membrane occurs at a rate that is 20 times slower than that through the polymer matrix, thus diffusion through the membrane is rate-limiting, which again improves the matrix-type square root of time-release kinetics, so that the release is like the zero-order release rate of a reservoir device. Following implantation in the upper arm, a single rod of Implanon releases 3-ketodesogestrel at the rate of > 30 µg/day for up to 3 years.

EVA copolymers are also used in fabricating Progestasert and Ocusert which are an intrauterine and an ocular drug delivery device for pilocarpine and progesterone, respectively. These are discussed in Chapters 11 and 12.

4.5

BIODEGRADABLE POLYMERIC IMPLANTS

Since the 1950s, most implants have been fabricated from nonbiodegradable, inert polymers such as silicone rubber, polyacrylamide and poly(ethylene-vinyl acetate) copolymers. However, some fundamental limitations of such implants include:

• The implants must be surgically removed after they are depleted of drug.

• Water-soluble or highly-ionized drugs and macromolecules, such as peptides and proteins, have negligible diffusivities through dense hydrophobic membranes.

• It is difficult to achieve versatile release rates—drug release rate is determined largely by the intrinsic properties of the polymers.

Such limitations prompted scientists to develop biodegradable polymeric implants. Degradation can take place via:

• bioerosion—the gradual dissolution of a polymer matrix;

• biodegradation—degradation of the polymer structure caused by chemical or enzymatic processes.

Degradation can take place by one or both mechanisms. For example, natural polymers such as albumin may be used; such proteins are not only water-soluble, but are readily degraded by specific enzymes. The terms degradation, dissolution and erosion are used interchangeably in this chapter, and the general process is referred to as polymer degradation.

Thus polymers used in biodegradable implants must be water-soluble and/or degradable in water. Table 4.4 lists some of the water soluble and biodegradable polymers that can be used for the fabrication of biodegradable implants.

Polymer degradation is classified into two patterns (Figure 4.11):

• bulk erosion;

• surface erosion.

In bulk erosion, the entire area of polymer matrix is subject to chemical or enzymatic reactions, thus erosion occurs homogeneously throughout the entire matrix Accordingly, the degradation pattern is sometimes termed homogeneous erosion.

In surface erosion, polymer degradation is limited to the surface of an implant exposed to a reaction medium. Erosion therefore starts at the exposed surface and works downwards, layer by layer. Due to the

Table 4.4 Synthetic polymers used in the fabrication of biodegradable implants

Water-soluble polymers	Degradable polymers

Poly(acrylic acid)	Poly(hydroxybutyrate)
Poly(ethylene glycol)	Poly(lactide-co-glycolide)
Poly(vinylpyrrolidone)	Polyanhydrides

difference in degradation rates between the surface and the center of the polymer matrix, the process is alternatively termed heterogeneous erosion. A drug distributed homogeneously in a surface-eroding matrix

Figure 4.11 Bulk and surface dissolution of biodegradable polymers

implant, of which the surface area is invariant with time, shows constant release with time over the period of implantation.

Polymer characteristics (type of monomer, degree of cross-linking, etc.) play a crucial role in determining whether the polymer is bulk- or surface-eroding. If water is readily able to penetrate the polymer, the entire domain of polymer matrix is easily hydrated and the polymer undergoes bulk erosion. On the contrary, if water penetration into its center is limited, the erosion front is restricted to the surface of the polymer matrix and the implant undergoes surface erosion. In practice, the polymer degradation occurs through a combination of the two processes.

As for non-degradable polymeric implants, biodegradable polymeric implants are divided into two main types:

• reservoir devices in which the drug is surrounded by a rate-controlling polymer membrane (such devices are particularly used for oral-controlled release—see Section 6.6.3);

• matrix devices in which the drug is distributed throughout the polymer matrix.

The drug release for biodegradable polymeric implants is governed not by diffusion through a membrane, but by degradation of the polymer membrane or matrix.

If the rate of polymer degradation is slow compared to the rate of drug diffusion, drug release mechanisms and kinetics obtained with a biodegradable implant are analogous to those provided by a nonbiodegradable implant (therefore a reservoir system gives a zero-order release profile and a matrix system gives a square root of time release profile). After drug depletion, the implant subsequently degrades at the site of implantation and eventually disappears.

However, in many cases, drug release takes place in parallel with polymer degradation. In such cases the mechanism of drug release is complicated as drug release occurs by drug diffusion, polymer degradation and/or polymer dissolution. The permeability of the drug through the polymer increases with time as the polymer matrix is gradually opened up by enzymatic/chemical cleavage. The references cited at the end of this chapter deal with the relevant mathematical treatments of this topic.

4.5.1

Poly-lactide and poly-lactide-co-glycolide polymers

Polyesters, such as poly(lactic acid) (PLA) and poly(lactic-co-glycolic acid) (PLGA), are examples of biomaterials that are degraded by homogeneous bulk erosion.

The polymers are prepared from lactide and glycolide, which are cyclic esters of lactic and glycolic acids. The lactic acid can be in either the L(+) or D(−) form, or the DL-lactic acid mixture can be used. Low molecular weight polymers (< 20,000 g/mol) are directly synthesized from lactic and glycolic acid via polycondensation. High molecular polymers (> 20,000 g/mol) are prepared via ring-opening polymerization (Figure 4.12). Variations in lactic acid:gycolic acid ratios, as well as molecular weights, affect the degree of crystallinity, hydrophobicity/hydrophilicity, and water uptake. Lactic acid-rich copolymers are more stable against hydrolysis than glycolic acid-rich copolymers.

Polymer degradation generally takes place in four major stages:

• Polymer hydration causes disruption of primary and secondary structures.

• Strength loss is caused by the rupture of ester linkages in the polymers.

• Loss of mass integrity results in initiation of absorption of polymeric fragments.

• Finally smaller polymeric fragments are phagocytosed, or complete dissolution into glycolic and lactic acids occurs (Figure 4.12).

4.5.1.1

Zoladex

Zoladex is a commercially available PLA/PLGA implant, designed to deliver goserelin (a GnRH agonist analog) over a 1- or 3-month period. As described in Chapter 1 (Section 1.5.2), chronic administration of GnRH agonists evokes an initial agonist phase, which subsequently causes antagonistic effects and a suppression of gonadotrophin secretion. Thus implants of GnRH analogues can be used clinically in the treatment of sex-hormone responsive tumors and endometriosis.

Zoladex implants are indicated for use in the palliative treatment of advanced breast cancer in pre- and peri-menopausal women, in the palliative treatment of advanced carcinoma of the prostate and in the management of endometriosis, including pain relief and the reduction of endometriotic lesions. The implant is fabricated by dispersing goserelin in a PLGA matrix and molding it into a cylindrical shape, which can be injected subcutaneously.

The release profile of goserelin from the implants has been well characterized during product development. For example, in a study of a Zoladex implant loaded with 10.8 mg of drug, the goserelin present at the surface of the implant was released rapidly, so that mean concentrations increased and reached peak levels within the first 24 hours. The initial release was then followed by a lag period up to 4 days, in which there was a rapid decline in the plasma concentration of the drug.

The lag period represents the time required to initiate polymer degradation. As water penetrates the polymer matrix and hydrolyzes the ester linkages, the essentially hydrophobic polymer becomes more hydrophilic. Extensive polymer degradation is followed by the development of pores or microchannels in the polymer matrix, which are visible by scanning electron microscopy (Figure 4.13). After the initial induction period required to initiate polymer degradation, drug release is accelerated thereafter by polymer degradation. In the above study this maintained the mean goserelin concentrations in the range of about 0.3 to 1 ng/ml until the end of the treatment period.

Figure 4.12 Synthesis and in vivo degradation of PLGA polymers

Figure 4.13 Scanning electron micrograph of a PLGA matrix incubated in distilled water at (37°C for 21 days). Pores and channels produced by extensive polymer degradation are visualized in the micrograph. The bar size is 1 µm. (Reproduced from Journal of Applied Polymer Science, 58: 197–206, 1995)

The discontinuous two phases drug release can be controlled and avoided by manipulating the degradation properties of the polymer so that it is possible for the Zoladex implant to provide continuous release over a 28-day period.

4.5.1.2

Lupron depot

The Lupron Depot comprises a PLA/PLGA microsphere delivery system for the delivery of the GnRH analog, leuprolide, over a 1-, 3-, or 4-month period. The release rate is determined by the polymer composition and molecular weight (Table 4.5).

The Lupron Depot microspheres are indicated for the treatment of male patients with prostate cancer and female patients suffering from endometriosis and anemia due to fibroids. Each depot formulation is

Figure 4.14 The chemical structure of poly[bis(p-carboxyphenoxy)propa ne: sebacic acid] and the pathway and products of its metabolism

Table 4.5 Lupron Depot characteristics

Release Rate	Polymer Composition	Polymer MW

1 month	PLGA (75:25)^a	12,000 to 14,000
3 or 4 months	PLA	12,000 to 18,000

^alactic acid:glycolic acid monomer ratio.

supplied in a single dose vial containing lyophilized microspheres and an ampoule containing a diluent. Just prior to intramuscular injection, the diluent is withdrawn by a syringe and injected into the single-dose vial to homogeneously disperse the microspheres.

An initial burst release of leuprolide from the microsphere depot occurs in vivo, followed by quasi-linear release for the rest of the time period. The efficacy of leuprolide depot formulations was found to be the same as the efficacy achieved with daily subcutaneous injections of 1 mg leuprolide formulation.

4.5.2

Polyanhydrides

Polyanhydrides, such as poly[bis(p-carboxyphenoxy)propane:sebacic acid] copolymers (Figure 4.14), are also used for the fabrication of biodegradable implants. Polymer degradation occurs via hydrolysis, the biscarboxyphenoxypropane monomer is excreted in the urine and the sebacic acid monomer is metabolized by the liver and is expired as carbon dioxide via the lung (Figure 4.14).

Erosion rates of poly (anhydride) copolymers are controlled by adjusting their molecular weight and biscarboxyphenoxy propane:sebacic acid ratio. Sebacic acid-rich copolymers display much faster degradation rates than biscarboxyphenoxy propane-rich copolymers. Changes in the ratio of the monomers are reported to provide various degradation rates ranging from 1 day to 3 years.

4.5.2.1

Gliadel

Gliadel is a biodegradable polyanhydride implant composed of poly[bis(p-carboxyphenoxy) propane:sebacic acid] in a 20:80 monomer ratio, for the delivery of carmustine. The implant is indicated in the treatment of recurrent glioblastoma multiforme (GBM) which is the most common and fatal type of brain cancer.

To fabricate the implant, the polyanhydride and the drug moiety are dissolved in dichloromethane. The solution is spray dried to produce microspherical powders in which the drug is homogeneously dispersed. The powders are then compressed into a disk-shaped wafer, approximately 14 mm in diameter and 1 mm thick.

Up to eight Gliadel wafers are implanted in the cavity created when a neurosurgeon removes the brain tumor. The wafers gradually degrade in the cavity and allows the delivery of high, localized doses of the anticancer agent for a long period, thereby minimizing systemic side-effects. Preliminary clinical reports with this system are highly encouraging.

In contrast to bulk-eroding PLA/PLGA polymers, the polyanhydride undergoes surface erosion. The thin-disk type morphology of the wafer confers a high surface-to-volume ratio on the implant, so that the total surface area of the implant is kept almost constant over the time of polymer degradation, which facilitates a constant release of carmustine with time.

4.5.3

Other biodegradable polymers

4.5.3.1

Poly(ortho esters)

Poly(ortho esters) offer the advantage of controlling the rate of hydrolysis of acid-labile linkages in the backbone by means of acidic or basic excipients physically incorporated in the matrix. This results in polymer degradation proceeding purely by surface erosion, which results in zero-order drug release from disk-shaped devices.

4.5.3.2

Poly(caprolactones)

Poly(-caprolactone) (PCL) is synthesized by anionic, cationic or coordination polymerization of ε-caprolactone. Degradable block copolymers with polyethylene glycol, diglycolide, substituted caprolactones and l-valerolactone can also be synthesized. Like the lactide polymers, PCL and its copolymers degrade both in vitro and in vivo by bulk hydrolysis, with the degradation rate affected by the size and shape of the device and additives.

Figure 4.15 Schematic illustration of the delivery of a drug to local tissues via the collagen implant injectable gel technology (courtesy of Matrix Pharmaceuticals Inc., Fremont, CA, USA)

4.5.3.3

Poly(hydroxybutyrate)

Poly(hydroxybutyrate) may be synthesized by fermentation from Alcaligenes eutrophus. The polymers have been shown to be useful for the controlled release of buserelin (a GnRH agonist analog) in both rats and humans.

4.5.4

Natural biodegradable polymeric implants

In addition to synthetic biodegradable polymers discussed so far, naturally occurring biopolymers have also been used for fabricating implantable drug delivery systems. Examples of natural biopolymers are proteins (e.g. albumin, casein, collagen, and gelatin) and polysaccharides (e.g. cellulose derivatives, chitin derivatives, dextran, hyaluronic acids, inulin, and starch).

Collagen, a major structural component of animal tissues, is being used increasingly in various biomedical and cosmetic applications. After implantation, collagen provokes minimal host inflammatory response or tissue reaction and its initial low antigenicity is practically abolished by the host’s enzymatic digestion.

A collagen-based therapeutic implantable gel technology has recently been developed, in which the drug moiety (a chemotherapeutic agent) is incorporated within the meshwork of rod-shaped collagen molecules. The collagen matrix is then converted to an injectable gel by a chemical modifier. Changes in the composition and structure of the gel can adjust its solubility, strength and resorption properties.

Direct injection of the gel into solid tumors and skin lesions provides high local concentrations of a drug specifically where needed (Figure 4.15). The gel is injected intradermally in a fanning or tracking manner to disperse the gel formulation throughout the tumor. Drug retention at the site of implantation is further enhanced by the addition of chemical modifiers such as the vasoconstrictor, epinephrine (adrenaline). This adjunct reduces blood flow and acts as a chemical tourniquet to hold the therapeutic agent in place.

The most advanced products based on the implantable gel technology include the Intradose (cisplatin/ epinephrine) injectable gel for treatment of solid tumors and the Advasite (fluorouracil/epinephrine)

Figure 4.16 Process of osmosis: the influx of water across a semipermeable membrane

injectable gel for treatment of cutaneous diseases, including external warts, basal cell carcinoma, squamous cell carcinoma and psoriasis.

4.6

IMPLANTABLE PUMPS

The driving force for drug release from a pump is a pressure difference that causes the bulk flow of a drug, or drug solution, from the device at a controlled rate. This is in contrast to the polymeric controlled release systems described above, where the driving force is due to the concentration difference of the drug between the formulation and the surrounding environment. Pressure differences in an implantable pump can be created by osmotic or mechanical action, as described below.

4.6.1

Osmotic implantable pumps

Osmosis is defined as the movement of water through a semi-permeable membrane into a solution. The semi-permeable membrane is such that only water molecules can move through it; the movement of solutes, including drugs, is restricted (although the extent of this restriction depends on the characteristics of the membrane, see below).

If a solution containing an osmotic agent (e.g. NaCl) is separated from water by a semipermeable membrane, the water will flow across the semipermeable membrane, into the solution containing the osmotic agent (Figure 4.16). Osmosis results in an increase in pressure in the solution and the excess pressure is known as the osmotic pressure.

The volume flow rate arising from the influx of water into the solution is determined by a number of factors:

• The osmotic pressure: ∆π is the difference in the osmotic pressure between osmotic agent-containing, and osmotic agent-free, compartments.

• The back pressure: water influx into the osmotic compartment generates a back pressure which retards the volume flow rate of water. ∆P is the difference in the hydrostatic pressure between the two compartments and represents the degree of back pressure generated.

• The effective surface area, A, of the membrane.

• The thickness of the membrane, h.

• The membrane selectivity toward an osmotic agent and water, described by the osmotic reflection coefficient σ. An ideal semipermeable membrane has the σ value of 1, which means that it allows the passage of only water molecules. In contrast, a leaky semipermeable membrane with a value approaching zero does not exhibit such selectivity and permits the transport of not only water, but also an osmotic agent.

• the permeability coefficient of the membrane, L_p.

These parameters affecting the volume influx of water can be expressed by:

(Equation 4.9)

Common semipermeable membranes and osmotic agents used in osmotic pumps are summarized in Table 4.6.

Osmotic pressure can be used for controlled drug release. The osmostic pressure can pump out drug at a constant rate, as described below. An important consideration is that because the pumping principle is based on osmosis, pumping rate is unaffected by changes in experimental conditions. Hence, in vitro drug release rate is often consistent with the in vivo release profile.

Table 4.6 Semipermeable membranes and osmotic agents commonly used in osmotic pressure-activated implantable pumps

Semi-permeable membrane

Cellulose acetate derivatives

Cellulose acetate, Plasticized cellulose triacetate, Cellulose acetate methyl carbamate, Cellulose acetate Ethyl- carbamate, Cellulose acetate phthalate, Cellulose acetate succinate

Other polymers

Poly(ethylene-vinyl acetate), Highly plasticized polyvinyl chloride,

Polyesters of acrylic acid and methacrylic acid, Polyvinylalkyl ethers,

Polymeric epoxide, Polystyrenes

Inorganic osmotic agents

Sodium chloride, Sodium carbonate, Sodium sulphate, Calcium sulphate, Mono- and Di-basic potassium phosphate, Magnesium chloride, Magnesium sulphate, Lithium chloride

Organic osmotic agents

Calcium lactate, Magnesium succinate, Tartaric acid, Acetamide, Choline chloride

Carbohydrates

Glucose, Lactose, Mannitol, Sorbitol, Sucrose

Swelling hydrogels

Sodium carbopol

Figure 4.17 Cross-sectional view of the Alzet osmotic pump, showing the various structural components

4.6.1.1

Alzet miniosmotic pumps

The Alzet miniosmotic pump consists of (Figure 4.17):

• Semipermeable membrane: serves as the housing for the entire pump and allows only water molecules to migrate into the osmotic sleeve.

• Osmotic chamber: contains sufficient contents of an osmotic agent.

• Reservoir wall: a cylindrical cavity molded from a synthetic elastomer, which is easily deformable by gentle squeeze. The fexible reservoir wall is impermeable to water molecules.

• Drug reservoir: contains the drug in solution/suspension.

• Flow moderator: a stainless steel, open-ended tube with a plastic end-cap, which serves as a pathway for the exit of drug solution/ suspension.

In this process, water crosses the outer semi-permeable membrane of the pump. The characteristics of the semipermeable membrane including permeability, pore size, and thickness are key factors determining the rate at which water molecules enter the osmotic sleeve. The water that is drawn across the semipermeable membrane causes the osmotic chamber to expand. This force compresses the flexible drug reservoir, discharging the drug solution through the flow moderator.

The osmotic pump can deliver a drug at a constant rate, if:

• the osmotic sleeve contains a sufficient amount of an osmotic agent so that the osmotic pressure remains unchanged for the duration of implantation, and

• the drug reservoir contains a saturated solution of the drug (this ensures that the concentration of drug is constant).

The selection of a semipermeable membrane is equally important since its properties, including A, h and σ, affect drug permeation (see Equation 4.9).

Alzet miniosmotic pumps permit easy manipulation of drug release rate (0.25 ~ 10 µl/hr) over a wide range of periods (1 day to 4 weeks). Also, as stated above, in vitro drug release rate from the osmotic pumps is often consistent with the in vivo release profile. These advantages mean that the miniosmotic pumps are used widely in experimental animal studies, to investigate, for example, the effects of drug administration regimen upon dose-response curve, as well as pharmacokinetic and pharmacodynamic profiles and drug toxicity. Alzet osmotic kits are also available, which allow the localized administration of drugs to the central nervous system of animals.

4.6.1.2

Duros implant pump

The Duros implant pump is a modified version of the Alzet miniosmotic pump which additionally contains a piston to control drug flow, between the osmotic engine and the drug resorvoir (Figure 4.18).

Water is drawn in across the semipermeable membrane and results in the expansion of the osmotic chamber. This force is delivered via the piston to the drug reservoir, forcing the contents of the drug reservoir to exit through the orifice.

Duros technology is demonstrating considerable promise for the controlled delivery of peptides and proteins. For example, a single implantation of the Duros pump in animals resulted in constant-rate release of biologically active leuprolide acetate (a GnRH analog) for one year. The use of non-aqueous vehicles to disperse peptides/proteins in the reservoir compartment is also being investigated. Although peptides and proteins are prone to degradation in aqueous solutions, adverse physicochemical reactions are sometimes avoided by dispersing them in a nonaqueous dispersion medium. Typical nonaqueous vehicles used in the drug reservoir compartment of the Duros implantable pump include waxes that soften around body temperature, hydrogenated vegetable oils such as peanut oil, sesame oil and olive oil, silicone oil, fatty acid monoglycerides or polyols. In addition, suspending agents, such as hydroxypropyl cellulose, poly(vinyl pyrrolidone) and poly(acrylic acid) are added to minimize the sedimentation rate of proteins inside the reservoir compartment.

4.6.2

Mechanical implantable pumps

The advance in microelectronics in the 1970s provided the momentum to develop externally programmable implantable pump systems. Such pumps were finally developed in the early 1980s and they allow physicians and patients to precisely control the infusion rate of a drug. Thus externally programmable pumps can facilitate

100

Figure 4.18 Cross-sectional view of the Duros implant, showing the various structural components

• zero-order drug release

• intermittent drug release.

Most implantable pumps are made of titanium which has proven records of excellent biocompatibility and long life. They are usually implanted intraperitoneally, in a pocket created in the abdominal wall of patients, under the subcutaneous fat layers, but above the muscular fascias. They are secured to the muscular fascia by suturing, which prevents pumps from flipping over or migrating in the pump pocket, thereby allowing patients to resume routine physical activities. Intraperitoneal insulin pump therapy is advantageous over a subcutaneous injection, as insulin infused into the peritoneal membrane surrounding abdominal organs is absorbed faster and more completely than via subcutaneous injection. Arterial or intraspinal delivery is also possible with a proper surgical procedure. A silicone rubber catheter is attached to the pumps, through which infusate is delivered to various body sites. The catheter is replaced if it becomes blocked, for example, by the deposition of infusate inside the lumen, fibrous tissue encapsulation or clotting at the tip.

101

4.6.2.1

SynchroMed implantable pump

The SynchroMed implantable pump was the first externally programmable implant pump to be introduced in the United States (in 1988). The major components are a miniature peristaltic pump, a drug reservoir (18 ml), a battery, an antenna, a microprocessor and a catheter through which infusate is delivered to a specific site.

The infusion rate of a drug solution can be programmed by a portable computer with specialized software which transmits instructions by radiotelemetry to the pump. The pump is driven by a step motor, controlled by signals from the micropocessor and is capable of delivering infusate at varying rates (0.004–0.9 ml/hr). The programmer provides the implantable pump with versatile delivery patterns, including a straight continuous-flow pattern or a more complex pattern that allows a varying dose at different times of the day to meet the patient’s changing metabolic requirements.

The SynchroMed pump is approved for use in:

• chemotherapy (using floxuridine, doxorubicin, cisplatin, or methotrexate);

• the treatment of chronic, intractable cancer pain (using morphine sulfate);

• osteomyelitis treatment (using clindamycin);

• spasticity therapy (using the muscle relaxant, baclofen).

However, the SynchroMed pump is not suitable for the delivery of insulin. The pressure of the roller heads on the tubing in the peristaltic pump causes intensive shear stresses which lead to stability problems for labile peptides and proteins.

4.6.2.2

MiniMed implantable pump

In the MiniMed implantable pump, a piston pump drives insulin through the delivery catheter. A patented solenoid motor controls the piston movement, to aspirate insulin from the reservoir chamber into the piston chamber and then push it through the insulin delivery catheter.

A hand-held programmer can change the pumping rate to administer the desired insulin dose to the diabetic patient. Thus the pump can be responsive to the diabetic’s fluctuating insulin needs. Many conventional insulin preparations are prone to denaturation when exposed to body fluids and temperature, or when agitated (see Section 1.6.1). The ensuing aggregation and precipitation may cause blockage of the catheter attached to the pump. However, the Minimed pump uses an insulin formulation, developed by Hoechst, which includes a small amount of Genapol (polyethylene glycol and polypropylene glycol), to increase the stability of the polypeptide.

4.6.2.3

Arrow implantable pump

The Arrow implantable pump is non-programmable and delivers infusate (2-deoxy-5-fluorouridine, morphine sulfate, baclofen, or heparinized saline) at 3 pre-set flow rates. The pump is divided into two chambers by accordion-like movable bellows. Infusate is placed in the inner drug reservoir chamber and Freon propellant in the outer chamber (Figure 4.19).

102

Figure 4.19 The cross-sectional view of the Arrow model 3000 implantable pump, showing the pumping mechanism

Drug delivery from this pump is powered by the Freon propellant. When the Arrow pump is implanted subcutaneously, it is warmed by the patient’s body temperature so that the propellant-containing chamber expands and exerts pressure on the movable bellows. Infusate is thus forced out of the reservoir chamber to an attached catheter through a filter and flow restrictor. This mechanism allows the delivery of infusate at a fairly constant rate to surrounding tissues or blood vessels. It should be noted, however, that the vapour pressure exerted by the outer chamber can be affected by changes in altitude/elevation or body temperature.

The Infusaid pump is another fixed-rate implantable pump that shares many similar features, including the Freon pumping principle, with the Arrow pump.

4.7

CONCLUSIONS

Implantable devices possess many advantages for drug delivery. Many different types of system are available and technology is expanding rapidly. Indeed, there now exists bio-responsive implantable systems, and implants for gene therapy; such advances are described in Chapter 16 (New Generation Technologies). However, despite the striking advances in this field, implantable systems will always be limited by the invasive nature of this therapy

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