Nanotechnological approaches for the delivery of macromolecules

Journal of Controlled Release 87 (2003) 81–88 www.elsevier.com / locate / jconrel

Nanotechnological approaches for the delivery of macromolecules Daan J.A. Crommelin a,b , *, Gert Storm a , Wim Jiskoot a , Robert Stenekes a,c , Enrico Mastrobattista a,d , Wim E. Hennink a a

Department Pharmaceutics, Utrecht Institute for Pharmaceutical Sciences ( UIPS), Utrecht University, P.O. Box 80082, Utrecht 3508 TB, The Netherlands b OctoPlus, Leiden, The Netherlands c Polymer Service Centre Groningen, Groningen, The Netherlands d MRC Laboratory of Molecular Biology, PNAC ( Biotech) Division, Cambridge, UK

Abstract In this overview, novel approaches are described for the controlled release and / or for the targeted delivery of macromolecules such as proteins and DNA. The building stones of these highly complex systems are (phospho)lipids and / or (biodegradable) polymers. They should be carefully chosen and preparation protocols should be rationally designed to maximize chances for success.  2002 Elsevier Science B.V. All rights reserved. Keywords: Drug targeting; Nanotechnology; Liposomes; Polymers; Polyplexes; Lipoplexes

1. Introduction Abbreviations: Dex, dextran; DS, degree of substitution; DTA, diphtheria toxin A chain; HEMA, hydroxyethyl methacrylate; hGH, human growth hormone; IgG, immunoglobulin; ILPP, immuno-lipopolyplexes; LCST, lower critical solution temperature; LPP, lipopolyplexes; MA, methacrylate; mAB, monoclonal antibody / bodies; PAA, poly(acrylic acid); PBT, poly(butylene terephthalate); PEG, poly(ethylene glycol); PEGT, poly(ethylene glycol)-terephthalate; PEO, poly(ethylene oxide); PL(G)A, poly(lactic(-co-glycolic) acid); PPO, poly(propylene oxide); PVA, poly(vinylalcohol); DPPC, dipalmitoylphosphatidylcholine; DPPG, dipalmitoylphosphatidylglycerol; PG, phosphatidylglycerol; pDMAEMA, poly-(2-(dimethylamino)ethyl methacrylate; OG, octylglucoside; chol, cholesterol *Corresponding author. Tel.: 131-30-2536-973; fax: 131-302517-839. E-mail address: [email protected] (D.J.A. Crommelin).

Macromolecules such as proteins and DNA play an increasingly important role in our arsenal of therapeutic agents. Delivery of these molecules to their site of action at the desired rate is a challenge because their transport through compartmental barriers (e.g., endothelium or epithelium) in the body is inefficient and / or they are readily metabolized. For controlled release or site-specific delivery of such macromolecules, delivery systems are required that are more sophisticated than our present delivery strategies. Those systems must be custom-made, taking into account both molecular size and specific characteristics of these molecules. One has to build a platform of different delivery strategies that use input

0168-3659 / 02 / $ – see front matter  2002 Elsevier Science B.V. All rights reserved. doi:10.1016 / S0168-3659(03)00014-2

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from technical, pharmaceutical and (bio)medical disciplines to meet this challenge. The rational design of those delivery systems requires an in depth insight of their structures at the molecular or ‘nanometer’ level. Two different strategies under investigation, one for controlled release of macromolecules and the other one for the targeted delivery of macromolecules, will be discussed in this concise review.

2. Custom-made carriers for controlled release of macromolecules Research on PLGA (poly(lactide-co-glycolide)) microspheres has dominated the field of parenteral peptide and protein sustained release systems in the last decade and led to successful commercial products. However, PLGA microsphere technology for proteins shows a number of problems [1,2]. The intimate contact of proteins with organic solvents in most preparation protocols and the interaction with this hydrophobic polymer after microsphere formation, in conjunction with the often-used high shear conditions during conventional manufacture proce-

dures often leads to (partial) inactivation of the proteins, incomplete release of encapsulated protein, and high initial burst in the release profiles. During release, a pH drop inside the spheres during microsphere degradation has been observed, which can further compromise the protein structure. These problems call for alternative matrix systems without these disadvantages. Microspheres based on hydrogels are now being developed as an alternative delivery system for PLGA spheres. Table 1 lists the proposed matrices that have been mentioned in the past [3]. If properly designed and manufactured, these hydrophilic structures may release proteins over prolonged periods of time with release kinetics dependent on, e.g., crosslinking strategies used. Existing cross-linking strategies for hydrogels were recently critically reviewed by Van Nostrum and Hennink [4]. Here we will focus on dextran as matrix material.

2.1. Dextran-hydrogel-based delivery systems Chemically and physically cross-linked dextran hydrogels were described before [5–7]. In the pro-

Table 1 Hydrogel matrices Based on natural materials: Collagen Gelatin Starch Alginates Chitosans Dextrans Based on synthetic polymers: Poly(N-vinylpyrrolidone) Poly(vinyl alcohol) Polyphosphazenes Poly(ethylene oxide)-b-poly(propylene oxide) copolymers (co-polymers, Pluronics) PL(G)A / PEO / PL(G)A co-polymers PVA-g-PLGA graft-polymers PEGT-PBT co-polymers (PolyActive) MA-oligolactide-PEO-oligolactide-MA Responsive polymers: Methacrylates (pH dependent swelling) Poly(N-isopropylacrylamide) (LCST) PEO-PPO-PEO (Pluronics) PEO-PPO-PAA graft-co-polymer (LCST) PLGA-PEO-PLGA (LCST) From Ref. [3].

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Fig. 1. Cumulative release of IgG from degrading dex-HEMA microspheres in time in vitro at pH 7, 37 8C. Water content 60%, DS 3 (♦) and water content 50%, DS 3 (j), DS 6 (d), DS 8 (m) and DS 11 (.). The values are the mean of two independent measurements that deviated less than 5% from each other. DS, degree of substitution, a measure for cross-linking density [8].

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cedures leading to the design of protein-loaded microspheres, the dextran-based technology showed remarkable characteristics. Protein release patterns could be varied over wide time ranges (days to months) (Fig. 1) [8] with very little burst release. This figure gives an example of microspheres designed in such a way that pulsed release patterns occurred, opening up opportunities to develop oneshot-vaccines with a built-in booster. The problem of interaction with organic solvents was elegantly solved by forming the dextran-microspheres in a PEG–water system. Microsphere formation is schematically shown in Fig. 2. There is no organic solvent (such as dichloromethane or chloroform) involved. The properties of the dextran microspheres depend on the chosen conditions. What parameters turned out to be critical in the formation of these dextran microspheres? PEG molecular weight and concentration, and dextran molecular weight, its

Fig. 2. Schematic representation of the microsphere preparation process for dextran microspheres. The protein is added to the aqueous phase and, in general, has a strong tendency to accumulate in the dextran microspheres, leading to high loading efficiencies. 10–15% of the dry microsphere material may consist of protein [3].

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Table 2 Length growth of PIT-1 dwarf mice model (no endogenous growth hormone production) upon injection of different doses of hGH-loaded dextran microspheres compared to daily injections of ‘free’ hGH (n55 mice) Length 2 weeks after injection / start of dosing (in cm); initial length 11.3 cm hGH-loaded microspheres (hGH content 166 mg), one injection

12.6

hGH-loaded microspheres (hGH content 83 mg), one injection

12.3

hGH, 10 injections of free hGH (total 83 mg) PBS control

12.35 11.6

degree of substitution by cross-linking groups plus its concentration are particularly important for the characteristics of the resulting microspheres. The structure of these microspheres and the protein were extensively studied in different states of hydration. Also, the mechanism of hydrogel degradation, protein release and protein integrity was elucidated. The thus-obtained new insights give ample opportunities to rationally fine-tune microsphere structure and release kinetics [9–12]. In vivo studies on biocompatibility gave positive results [13] and sustained therapeutic activity was observed in mouse experiments with hGH (human growth hormone) loaded microspheres (Table 2). In

this experiment, daily injections of hGH resulted in comparable growth rates as one single injection of hGH microspheres (same total hGH dose). This finding might lead to more patient friendly dosing schemes. An interesting finding was that colloidal particles, such as liposomes, could be efficiently loaded in these microspheres as well. They were released in intact form in a controllable manner. Here again, pulsed release patterns can be readily obtained with these ‘liposomes-in-microspheres’ systems (Fig. 3) [14]. This offers possibilities for the design of controlled release systems for liposomeencapsulated drugs, including macromolecules (e.g., membrane proteins).

3. Custom-made carriers for targeted delivery of macromolecules

Fig. 3. The release of DPPC–DPPG–chol (10:1:10, mol ratio) liposomes from dex-HEMA microspheres at pH 7.2 using different formulations: (d) DS 5 and 70% water; (m) DS 5 and 50% water; (.) DS 8 and 70% water; (j) DS 8 and 50% water and the release of liposomes from dex(lactate) 2 HEMA (♦) DS 4 and 70% water. Values are the mean of at least two measurements [14].

Targeted delivery of proteins and DNA requires a carrier system in the sub-micron size range. This carrier should be target site (cell, tissue) specific. Often the actual target site is intracellularly located and the delivery of the carrier-payload at this intracellular target site is a prerequisite for therapeutic success. For example, plasmid DNA has to be delivered inside the nucleus of the target cell before the cell can express the desired therapeutic protein. Two examples of targeted delivery strategies where we design tailor-made carriers will be discussed.

3.1. Delivery of ‘ liposome-dependent’ drugs In the first example, a ‘liposome-dependent’ drug

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is used. ‘Liposome-dependent’ means that the drug as such cannot reach its target site without a carrier. For example, its site of action is inside the cell and it cannot pass the cytoplasmic membrane without ‘help’ (a carrier). Such a drug will neither show the desired, nor the undesired pharmacological effect at any place in the body. An example of a liposome-dependent drug is diphtheria toxin A chain (DTA). Diphtheria toxin (a protein consisting of an A chain coupled to a B chain) can readily enter cells through the transporter B chain. Upon entering the cell, the A chain causes the cell kill with exceptional efficiency by blocking ribosomal activity. Thus, DTA (lacking the B chain) alone needs a cell-specific transport system, i.e., a system that transports it into the desired target cells, e.g., tumor cells. Earlier, we showed a high target cell (ovarian tumor cells) specific binding of intraperitoneally injected immunoliposomes in a mouse model with ovarian cancer cells growing in the peritoneal cavity [15] (Fig. 4). Immunoliposomes consist of liposomes carrying the drug with monoclonal antibodies or monoclonal antibody fragments

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covalently attached to the bilayer for targeting purposes. However, our first attempts to deliver these targeted immunoliposomes into the ovarian carcinoma target cells failed. Not all monoclonal antibodies (or their fragments) induce endocytic uptake after binding to target cell-specific receptor sites. Only after selecting a monoclonal antibody molecule that indeed induced endocytic uptake, the immunoliposomes could enter the tumor cell. The next step in the series of events to bring DTA into the cytosol was then to facilitate endosomal escape of the DTA liposome into the cytosol and escape of DTA from the liposome. To this end, a pH-dependent fusogen (influenza virus-derived peptide: diINF-7) was added to the DTA-carrying immunoliposome. Here, we have an excellent example of a drug carrier system in the nanometer size range that is rationally designed to efficiently deliver its payload at the target site (Fig. 5). The in vitro tests using an ovarian carcinoma cell line (OVCAR-3) demonstrated specific killing of OVCAR-3 cells. Controls such as free DTA, DTA in immunoliposomes with an irrelevant antibody or with no antibody at all did not show

Fig. 4. Immunoliposomes binding to the surface of an ovarian carcinoma cell. This electron micrograph depicts a human OVCAR-3 cell taken from the peritoneal cavity of nu / nu mice after injecting the animals intraperitoneally with OV-TL3-Fab’-immunoliposomes. A more detailed analysis of the cell-immunoliposome interaction showed very little endocytic uptake. A search was started to identify endocytosis inducing antibodies. mAB with human ovarian cancer cell specificity were identified (e.g., mAB 425). These 425 immunoliposomes loaded with DTA and a pH-dependent fusogen (diINF-7) were tested in vitro [15,17].

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3.2. Building a non-viral plasmid transfection system

Fig. 5. The use of nanotechnology to build a carrier system for site-specific delivery of a protein. The anti-EGFR antibody permits endocytosis helping the immunoliposomes to enter the target cells after cell binding. The peptide diINF-7 induces release of liposome-entrapped diphtheria toxin A chain (DTA; the ‘liposome dependent drug’) from the immunoliposomes present in the endosome / lysosomal compartments. Di-INF-7 expresses membrane destabilizing activities only under acid conditions as met in the endosome and thereby helps the DTA to enter the cytoplasm [17].

tumor cell killing [16]. In vivo evaluation of this concept in the nu / nu mouse model is presently in progress.

Another example of the design of a custom-made carrier at the nanometer level is a targeted delivery system for plasmid DNA in order to efficiently transfect only target cells [17]. The cationic polymer pDMAEMA (poly-(2-(dimethylamino)ethyl methacrylate)) condenses plasmid DNA effectively into 100-nm nanoparticles (polyplexes). In vitro transfection is rather efficient. However, in vivo transfection is hampered, probably because of early destabilization of the polyplex (before being taken up by the target cell) by negatively charged macromolecules that compete with the plasmid for interactions with the pDMAEMA. The polyplex then falls apart and hardly any transfection is observed. In vitro, this destabilizing effect can be monitored by adding the negatively charged biopolymer hyaluronic acid to the polyplex dispersions. Because of this early destabilization of the polyplexes we decided to coat them with lipids yielding so-called lipopolyplexes. The strategy for lipid coating is depicted in Fig. 6 [16]. Basically, the polyplexes are added to a mixed micellar solution of lipids and the detergent octylglucoside (OG). Lipid mixtures with a substantial

Fig. 6. Schematic representation of lipid-coated polyplex formation. Full details in Refs. [16,17]. (1,2) Plasmid DNA is condensed by adding the cationic polymer pDMAEMA to the DNA at a DNA / pDMAEMA ratio of 1:3 (w / w). (2,3) The formed lipoplexes are added to mixed micellar solutions containing the detergent octylglucoside (OG) and a total amount of 3 mmol of detergent-solubilized lipids. (3,4) Slow removal of OG by adsorption to hydrophobic BioBeads and formation of lipid coats preferentially around positively charged complexes because of electrostatic interactions.

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fraction of the negatively charged phospholipid, phosphatidylglycerol (PG), turned out to efficiently coat polyplexes and maintain an average particle size of around 100 nm. The positive zeta-potential on the original polyplexes was turned into a negative one. Indeed, hyaluronic acid did not destabilize these lipopolyplexes. The positive charge on the original polyplexes usually enables intense contact with any cell membrane encountered. Target cell specific transfection is then difficult to realize. With the lipopolyplexes the positive charge is gone and target cell specificity could be introduced by attaching specific monoclonal antibody fragments (Fab9 fragments of mAB 323 / A3) to the phospholipid coat. These immunolipopolyplexes showed target cell-specific binding characteristics (to OVCAR-3 cells). In addition, they were able to transfect the cells with hardly any cell toxicity (Fig. 7). In an attempt to further improve transfection efficiency and selectivity we are now working on an extension of this nanotechnological approach. Fig. 8 schematically depicts the concept (Fig. 8). Here, a combination of liposome technology and polymer technology is used. PEG-coated lipid structures

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surrounding cationic polymer condensed plasmid DNA are combined with an effective fusogen, nuclear localization signal and / or an antibody as homing device. This forms a sort of artificial virus that should both be site-specific and transfectionefficient. A question to be addressed, of course, is whether these ‘artificial viruses’ are indeed less immunogenic and more safe than their ‘real-viral’ counterparts, such as adenoviruses.

4. Conclusion Modern drug delivery systems can be custommade. One specific delivery system for the treatment of each disease. But, that implies that an in depth knowledge of the desired characteristics of these delivery systems is available. One should have a full insight in the physiological and pathological conditions relevant to achieve the desired drug action, their pharmacological properties and, last but not least, pharmaceutical issues such as stability, during preparation and on the shelf, chemical reactions involved, up-scaling possibilities and choice of excipients.

Fig. 7. Transfection and cell viability of OVCAR-3 cells after long-term exposure (48 h) to pDMAEMA-polyplexes, lipopolyplexes (LPP) and immuno-lipopolyplexes (ILPP). OVCAR-3 cells (1.1310 4 cells / well) were exposed for 48 h at 37 8C to either polyplexes, lipopolyplexes or immuno-lipopolyplexes (at 1 mg plasmid DNA / well) and subsequently evaluated for b-galactosidase expression and cell viability using the ONPG and MTT assay, respectively [16,17].

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Fig. 8. The ‘artificial virus’ approach.

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