Which polymers can make nanoparticulate drug carriers

advanced

drug deliiry reviews

ELSEVIER Advanced Drug Delivery Reviews 16 (1995) 141-155

Which polymers can make nanoparticulate drug carriers long-circulating?

Vladimir P. Torchilin”, Vladimir S. Trubetskoy Center for Imaging and Pharmaceutical Research, Department of Radiology, Massachusetts General Hospital and

Harvard Medical School, 149, 13th Street. Charlestown, MA 02129, USA

Accepted 13 April 1995

Abstract

The protective effect of poly(ethylene glycol) and some other polymers on nanoparticulate carriers including liposomes is considered in terms of statistical behavior of macromolecules in solution, when polymer flexibility plays a key role. According to the mechanism proposed, surface-grafted chains of flexible and hydrophilic polymers form dense “conformational clouds” preventing other macromolecules from the interaction with the surface even at low concentration of protecting polymer. Using liposomes as an example, experimental evidence is presented of the importance of protecting polymer flexibility in liposome steric protection. Further possible applications of the suggested model are discussed. The possibility of using protecting polymers other than poly(ethylene glycol) is analyzed, and examples of such polymers are given based on polymer-coated liposome biodistribution data. General requirements for protecting polymers are formulated, and differences in steric protection of liposomes and particles are discussed. The scale of protective effect is interpreted as the balance between the energy of hydrophobic anchor interaction with the liposome membrane core or with the particle surface and the energy of polymer chain free motion in solution.

Keywords: Long-circulating drug carriers; Liposomes; Nanoparticles; Poly(ethylene glycol); Amphiphilic flexible polymers; Polymer conformations

Contents

1. Introduction ..............................................................

2. Liposomes as model carriers to study the phenomenon of prolonged circulation

3. New polymers for steric protection of liposomes ............................. 4. Protective polymers on the surface of nanoparticles and latexes ............... 5. Conclusion.. .............................................................

References ..................................................................

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* Corresponding author. Fax: +1 617 7267830.

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0169-409X/95/$29.00 @ 1995 Elsevier Science B.V. All rights reserved

SSDI 0169-409X(95)00022-4

1. Introduction

Nanospheres, nanocapsules, liposomes, mi- celles, and other nanoparticulates are frequently referred as carriers for delivery of therapeutic and diagnostic agents [l-3]. Surface modification of these carriers is often used in order to increase longevity and stability of nanocarriers in the circulation, to change their biodistribution, and / or to achieve targeting effect. Recent trends involve the increasing use of different polymers for carrier surface modification. Synthetic or natural polymers have been shown to protect solid particulates from interaction with different solutes. The phenomenon is closely connected with stability of various aqueous dispersions and has a tremendous practical significance [4], in particular within the pharmaceutical field where the polymers might protect particulate drug carriers from undesirable interactions with biological milieu components.

In recent years, amphipathic polymers com- posed of at least two different parts in their molecule with pronounced hydrophobic and hydrophilic properties have gained increasing at- tention. On one hand, these polymers demonstrate the ability to be easily adsorbed (or incorporated) on the surface of the particulate carrier due to hydrophobic interactions; on the other hand, their hydrophilic portions are exposed to the solution and effectively protect those particulates from interactions with plasma proteins in the blood upon intravenous administration [5]. Another way to protect the surface of nanoparticles or latex particles with the polymer is to graft the polymeric chain onto the surface by covalent bond to form a so-called “anchored” polymer. The term “steric stabilization” has been introduced to describe such phenomena [6].

Independent of the carrier type and designa- tion, its modification with synthetic polymers results in the appearance of interesting theoretical and practical problems connected with the accessibility of the carrier surface for carrier- interacting substances. The most general ques- tion is: what are the peculiar properties of certain polymers which provide them with the ability to serve as effective steric stabilizers or steric protectors?

2. Liposomes as model carriers to study the phenomenon of prolonged circulation

Liposomes may serve as a good model for the understanding of grafted polymer influence on carrier properties, and many regularities found for liposomes might be successfully applied to many microparticulate drug carriers. One of the most popular and successful methods to obtain long-circulating biologically stable liposomes is their coating with certain hydrophilic and flexible polymers, first of all with poly(ethylene glycol) (PEG) [7-91. To make PEG capable of incorporation into the liposomal membrane, the reactive derivative of hydrophilic PEG is single terminus- modified with hydrophobic moiety (usually, the residue of phosphatidyl ethanolamine or long chain fatty acid is attached to PEG-hydroxysuccinimide ester, PEG-OSu) [7], see Fig. 1.

Starting from the first description of such liposomes [7], the mechanism of PEG protective action is under continuous investigation [lo-141. With the understanding of this mechanism some other polymers were expected to be suggested for liposome protection in vivo, thus widening the possible areas of biomedical application of liposomes. There are already several reviews, for example [15,16], describing different properties of long-circulating PEG liposomes including the mechanism of PEG protective action. The general opinion is that, on the biological level, coating

e F wr(-CH2-CH2-O-),, - ‘-(Cll&- -0-N

PEG-OSu

2

PE

Fig. 1. Synthesis of PEG-PE from PEG-hydroxysuccinimide

ester (PEG-OSu) and phosphatidyl ethanolamine (PE) [7].

142 V.P. Torchilin, VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-IS5

V.P. Torchilin, VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155 143

liposomes with PEG sterically hinders interactions of blood components with the liposome surface. This slows down the inhibition of liposome destruction by lipoproteins, and prevents liposome interaction with opsonins resulting in fast capture of liposomes by the reticuloendothelial system (RES). To confirm this, the reduced binding of plasma proteins with PEG liposomes was demonstrated in [10,15,17,18]. The role of the supposed inhibition of opsonization of PEG liposomes in the prolongation of their circulation time also seems evident [19].

Less evident is the mechanism by which PEG (or similar polymers) prevents opsonization and interaction of liposomes with proteins. The whole set of probable (possibly mutually addi- tive) mechanisms has been discussed in the literature. Thus, from the properties of colloids it is known that repulsive interactions between colloidal particles can be enhanced by coating these particles with soluble, well hydrated, and chemically inert polymers [6]. Such modification might decrease surface hydrophobicity and interaction of particles with RES. Coating of different nanoparticles with amphiphilic polymers, when hydrophobic block serves as an anchor to hx the polymer on the surface and hydrophilic block forms the protective layer around the particle, was shown to drastically decrease particle capture by RES and strongly influence particle biodistribution; many experiments were performed with particles coated with poloxamers [20-221. Special importance was also prescribed to the role of surface charge and hydrophilicity of PEG-coated liposomes [ 111.

However, surface hydrophilicity and repulsive interactions alone cannot explain all the phenomena observed. This was also noticed by Allen in her recent review [15]. Thus, for example, liposomes coated with well soluble and highly hydrophilic dextran did not demonstrate prolonged circulation time or ability to protect the liposome surface from interacting with proteins from the solution [14,23]. (In [24] dextran chains, however, were shown to protect sterically certain soluble molecules.) To understand further what peculiarities of the behavior of PEG and similar polymers underlie their ability to prevent liposome opsonization and blood clearance, we hy-

pothesized that the important feature of protective polymers is their flexibility (free rotation of individual polymer units around inter-unit link- ages). It means that the molecular mechanism of polymer protective action is determined by the properties of a flexible polymer in solution and includes the formation of the polymeric layer over the liposome surface which is impermeable for other solutes even at relatively low polymer concentrations [23,25-271. Evidently, the protective layer of a polymer over the liposome surface has to combine abilities to escape recognition by cells (the best way to be invisible is to look like water), and to prevent the penetration of opsonizing proteins to the liposome surface (which requires the polymer coat to be dense enough). We will discuss furthermore how these seemingly inconsistent properties can be combined. For simplicity, we assume that transient contacts between protective polymers and plasma proteins do not result in opsonization, and polymers themselves do not interact with cells.

Considering the diffusional movement of a macromolecule from solution towards the liposome surface as the initial step of its interaction with liposome, we can express the degree of liposome protection as the probability (P) for the protein to collide with a polymer instead of a liposome [27]. The higher the P value, the better the protection. When P equals 1, no interaction between protein and liposome surface is possible. To forecast possible probability values, we described the behavior of liposome-grafted polymer in terms of a simplified statistical model of a polymer solution [28]. Within this model, polymer solution is considered as a three-dimensional network, where each cell may be occupied either with a polymer unit or with a solvent (water) molecule. The more flexible the polymer, i.e., the more independent the motion of any polymeric unit relative to the neighboring one, the larger the total number of its possible conformations and the higher the transition rate from one conformation to another. As a result, water-soluble flexible polymer statistically exists as a distribution (“cloud”) of probable conformations. The polymer flexibility correlates with its ability to occupy with high frequency many cells in solution, temporarily squeezing water molecules

144 V.P. Torchilin, KS. Truberskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155

out of them and making them impermeable for other solutes which require free water for diffusion. Thus, a relatively small number of water- soluble and very flexible polymer molecules can create sufficient density of conformational “clouds” over the liposome surface and protect the latter from destruction or opsonization. Kinetic aspects of this model were discussed in [27]. From the kinetic analysis, we can easily obtain the following equation:

where y is the molar ratio polymer/lipid in the outer monolayer, S,,, the effective square protected by a single polymer molecule, S, the average area occupied by a single lipid molecule (for the given liposome size and composition), and P* the “reliability” parameter or the average P within the “cloud” volume. At high y values, polymer will be “stretched” out of the liposome. Therefore, y(S,IS,) is always ~1. The maximal protection can be achieved when yS, = S, (the polymer “clouds” are practically fused) and the P* value is close to 1.

In a rigid chain polymer unit motion is hin- dered, and even good water solubility and hydrophilicity may not provide sufficient protection for the liposome surface, as was noticed for liposome-grafted dextran [14,23]. The number of possible conformations for such polymers is lower, and transitions from one conformation to another proceed at a slower rate than for flexible polymer. The density of the conformational “cloud” for a rigid polymer and the number of water molecules disturbed will be much smaller. Sufficient free water space within the “cloud” will make the diffusion of plasma proteins to- ward the liposome surface still possible. So, to serve as an effective liposome protector even at relatively low concentration of the surface-immobilized macromolecules, the polymer has to combine hydrophilicity (solubility) and flexibility (which is characteristic of, for example, PEG).

The size of the area on the liposome surface protected with a single polymer molecule of a given molecular weight, and the number of polymer molecules required for the effective protection of the liposome of a given size were

calculated by us using the average end-to-end distance of a polymer random coil in solution parameter, R, ,,sO, [27]. Using &so, values for PEG of different molecular weights published in [29], we estimated the area of the liposome surface which can be protected by a single PEG molecule, and the molar ratio PEG-to-lipid required for complete protection of the liposome surface, which match well with published experimental data [9,30].

To develop the model of flexible polymer behavior on the particle surface, we applied computer simulation [23]. Spatial distribution of liposome surface-grafted polymers was simulated using a three-dimensional random flight model, where polymer was assumed to consist of abso- lutely rigid segments with free segment rotation around intersegmental conjunction (conditionally assuming the segment length I= 1 nm for a flexible polymer and I = 5 nm for a rigid one). From the computer analysis it follows (the cross- section of calculated protective “cloud” is presented in Fig. 2) that the flexible polymer forms the high density conformational cloud. A rigid polymer of the same length forms a broad but loose and thus permeable cloud.

An important consequence of the suggested model is that if there exist any reactive centers on the liposome surface (e.g., binding sites for certain macromolecules, antibodies, antigens, or other ligands), the presence of PEG in concentrations below saturating leads to the appearance of two populations of those centers. The first population will consist of reactive moieties excluded from the PEG-occupied volume. Such reactive centers should possess the same properties and reactivity as on “plain” liposomes. However, centers remaining within the polymer cloud form the second population with a sharply decreased ability to participate in “normal” interactions [23].

More detailed analysis within the frame of the suggested model permits to answer several important questions (V. Torchilin and B. Hoop, in preparation). For example, how are the PEG coils distributed in the space above the surface of the liposome? The thickness r of the PEG protective surface (the extent to which the PEG chain sticks up above the particle surface) can be

I/P. Torchiiin, KS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155 145

2, nm

15 . . I I

Fig. 2. Distribution of polymer conformations in space; slice X = 0 t 0.25 nm. Produced by random flight simulation (Z > 0, polymer length 20 nm, 440 conformations). Upper panel: segment length is 5 nm (rigid polymer); lower panel: segment length is 1 nm (flexible polymer). Calculations for the rigid polymer have been done under the assumption that only every fifth 1 nm segment may change the direction. The “rings” in the upper panel are the result of the assumption that the net mass is concentrated at the remote termini of “frozen” shorter segments. From [23].

-10 -5 0 5 10

Y, nm

represented by a class of probability distributions c(t) (chi-squared distributions):

4) = z-(p) 7

cop p’ ‘“-“e-(!L) ( > where c,, is a normalizing factor, r is the mean length of the PEG coil above the particle (liposome) surface, and p is the “shape” factor which determines the width of the distribution. The width of the distribution is given by p-’ and T(p) is the corresponding gamma function. For a value of p = 0.01, the distribution peaks at the surface, whereas for a large p (=20) the distribution approaches a gaussian shape centered about a normalized mean length t/r = 1. That is, as p goes from near zero to very large, the probability distribution goes from hyperbolic (p = 0.01; the PEG molecule is most likely lying

on the surface) to approximately gaussian (p = 20; the PEG molecule is most likely stretched out to its full length perpendicular to the liposome surface). The effective volume (bare liposome plus PEG-coated sheath) would then be just V= 4/37r(R + (T))~, where R is the bare liposome radius and (r) is the mean sheath thickness (mean length r weighted by c(r)).

This is just a single example of the possible application of the approach. Other important questions which can be analyzed within the suggested model are: (a) How does the space available for water molecules depend on the distance above the liposomal surface at different sizes and concentrations of surface-grafted PEG molecules?, and (b) What is the thickness of the spherical surface volume around a liposome which can be entirely filled with PEG molecules [31]? The answers for these questions might contribute to a further understanding of the mechanism of polymer-mediated liposome (particle) protection and development of optimized long-circulating microparticular drug carriers.

Our model was confirmed experimentally by studying the efficacy of liposome-incorporated fluorescent marker quenching by macromolecular quencher from the solution, depending on the polymer presence on the liposomal membrane [23]. Two different systems were investigated: (a) the quenching of liposome-incorporated N-[7- nitrobenz-Zoxa-1,3-diazol4-ylldioleoyl-phosphatidyl ethanolamine with soluble rhodamine- modified ovalbumin, and (b) the quenching of liposome incorporated phospholipid derivative of fluorescein with anti-fluorescein antibody. In both cases, the presence of even small quantities of flexible PEG (as low as 0.2 mol%) on the liposome sharply decreased the quenching compared with that for plain or rigid dextran-modified liposomes. Since the whole quenching process is limited only by macromolecular quencher diffusion from the solution to the liposome surface, it is evident that the presence of PEG on the surface creates pronounced diffusional hin- drances for this process [23]. Besides, the exist- ence of two different fluorescein pools on the liposome surface was revealed at low PEG concentration. One of them was quenched with the same rate as fluorescein on plain liposomes,

146 VP. Torchilin, VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155

whereas the quenching kinetics for another was close to that for fluorescein on PEG liposomes with higher PEG content. This reflects a different location of fluorescein molecules on the liposome surface-between and inside PEG “clouds”, as was predicted by theoretical analysis.

3. New polymers for steric protection of liposomes

The suggested model and experimental data (partially discussed above) permitted us to formulate some general requirements towards polymers for liposome protection. These polymers should be soluble and hydrophilic, and have a highly flexible main chain. Synthetic polymers of the vinyl series, such as poly(acry1 amide) (PAA) and poly(viny1 pyrrolidone) (PVP), may serve as the most evident examples of other potentially protective polymers. (The list of possible sterically protecting polymers for liposomes described so far is presented in Fig. 3).

We reported recently the first experimental data on the possible use of synthetic polymers other than PEG for the steric protection of liposomes in vivo [32,33]. Both amphiphilic PAA and PVP used in our study were prepared by radical polymerization of monomers in organic solvent in the presence of different quantities of long chain fatty acid chloroanhydride (used as a chain terminator introducing a terminal long chain acyl group into the polymer molecule). Both polymers were prepared with MW 6000- 8000 (low-molecular-weight polymers, PAA-L and PVP-L) and 12000-15000 Da (high-molecular-weight polymers, PAA-H and PVP-H). As a terminal hydrophobic anchor these polymers contained either a palmityl (P) or a dodecyl (D) group.

Biodistribution experiments in mice were performed with lecithin/cholesterol liposomes prepared by the detergent dialysis method with the addition of 2.5 or 6.5 mol% of the corresponding amphiphilic polymer. For clearance and biodistribution studies, liposomes (165-190 nm) were labeled with ’ ’ ‘In via membrane-incorporated diethylene triamine pentaacetic acid- stearylamine [9] prepared as described in [34].

It was shown that amphiphilic derivatives of PAA and PVP provide effective protection to liposomes in vivo similar to PEG-PE [7,23]. The extent of protective activity for different polymers depends on the length of hydrophobic anchor, polymer molecular weight (i.e., chain length and the energy of molecular motion), and the quantity of protecting polymer on the liposome surface. When modified with the same palmityl residue, PAA, PVP, and PEG of similar molecular weights (6000-8000) being used in similar concentrations, all provide efficient steric protection for liposomes and noticeably increase the residence time of liposomes in the blood. Half-clearance times for PVP-L-P, PAA-L-P, and PEG liposomes with 2.5 mol% content of protective polymer are ca. 45, 80, and 80 min, and for PVP-L-P, PAA-L-P, and PEG liposomes with 6.5 mol% content of protective polymer ca. 120, 140 and 170 min, respectively, whereas the half-clearance time for “plain” liposomes of the same size is only about lo-15 min.

The protective activity of polymers with shorter hydrophobic moiety or with higher MW (PAA-L-D, PAA-H-P and PVP-H-P) is, however, much lower (corresponding half-clearance times do not exceed 40 min). Despite some increase in the circulation time and decrease in the liver capture of liposomes, these polymers are far less effective steric protectors than polymers of the same molecular weight, but with longer acyl anchor (compare PAA-L-D and PAA-L-P liposomes), or polymers with the same long acyl anchor, but with higher molecular weight of the hydrophilic moiety (compare PAA- H-P and PVP-H-P liposomes with PAA-L-P, PVP-L-P, and PEG liposomes). This can be easily understood taking into account the supposed energy of interaction between the fatty acyl anchoring the polymer into the liposome surface and the hydrophobic part of the liposomal membrane. From the thermodynamic point of view, a relatively short dodecyl group is unable to keep a 6-8 kDa polymer molecule on the liposome surface: the energy of the polymeric chain motion in the solution might be higher than the energy of the dodecyl group interaction with phospholipid surroundings within the liposomal membrane. As a result, PAA-L-D

VP. Torch&, VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155

NH-(CH,),-O-k0

d

PEG-P& [71

mPEG -oCO-NH *

mPEG -OCO-NH

branched PEG-PI?, (351

H&Z- -N.C,,,-CH, -CO-(CH$,-CO- NH-(CH&-O-P.0

I 1 - J

6 R’ *o

d

n ,R=CH, , C,H,

poly(Z-alkyl-2-oxazoline)-PE, [381

H. :CH,-CH- I C’O

A

0 0

* -9 J== -S-CH,-CO- NH-(CH,),-O-P.0

8

n

poly(acryloy1 morpholine)-PE. [361

poly(wnyl pyrrolidone).PE. [371 ”

HO-lCH&C-CH2-0, ,I.OF

poly(glycerol)-phosphatidyl glycerol. [391

-CH,-CH-COW

poly(vinyl pyrmlidone)-palmitate. [32]

H- -CH,GH -

1 1 -CH&H -CO-

AONH, AONH, n

poly(acryl amide)-palmitate, [32l

147

Fig. 3. Amphiphilic synthetic polymers used for steric protection of liposomes. *Branched PEG-PE, poly(acryloy1 morpholine)-PE,

and poly(vinyl pyrrolidone)-PE were prepared by modification of corresponding carboxyl-terminated polymers according to

Torchilin and Trubetskoy, submitted.

148 V.P. Torchilin, KS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155

might be relatively easily removed from the liposomal membrane, and demonstrates only a slight and transient protective effect. The longer palmityl anchor provides much more firm polymer binding with liposome (higher energy of interaction with the hydrophobic membrane core due to the larger number of membrane-embed- ded CH, groups), and thus much better liposome steric protection. On the other hand, even the length of palmityl anchor is insufficient to fix firmly a 12-15 kDa polymer on the liposome surface because of much higher energy of polymer chain motion in the solution compared with that for the shorter polymers [28]. So, the liposome surface gradually looses the protective polymer coat; as a result the liposome is opsonized and captured by the liver and spleen.

Other amphiphilic polymers with well soluble and flexible hydrophilic moiety, such as amphiphilic poly(acryloy1 morpholine) (PAcM), branched PEG in which two PEG chains are attached to a single hydrophobic phospholipid group (PEG2) and phospholipid-attached PVP, have also been successfully used as liposome steric protectors in vivo (VP. Torchilin et al., J. Pharm. Sci., in press). Amphiphilic PEG2, PAcM, and PVP can be prepared in two steps. In the first step carboxyl-terminated polymers were synthesized (PAcM-COOH, PEG2-COOH, PVP- COOH, see their structures in Fig. 4). In the second step, carboxyl-terminated polymers were modified with PE according to the scheme presented in Fig. 5.

A branched carboxyl-terminated derivative of monomethoxypoly(ethylene glycol) (Fig. 4A), where two linear PEG chains are attached to a single COOH group, was prepared by activation of a terminal hydroxy group of monomethoxy- PEG(MW 5000) with 4-nitrophenyl chloroformate, subsequent coupling of lysine via the c-amino group, and further attachment of another 4-nitrophenyl chloroformate-activated monomethoxy-PEG to the free lysine a-amino group as described in (351. Alternatively, PEG2 was prepared by the direct interaction of lysine with the excess of PEG-succinimidyl carbonate [35]. To prepare PAcM-COOH (Fig. 4B), the corresponding monomer was synthesized by acylation of morpholine with acryloyl chloride,

mPEG -oco-NH

A

mPEG -OCO-NH /

H- -CH,-CH- -S-CH,-COOH r ,i

B +=0 N I J 0 0 n

0

H- r-CH,-_CH-l-C(CH,),-0-CH,-CH,-0-~.NH-CH,-COOH

C I: I I c-l 0

n

Fig. 4. Chemical structures of carboxyl-terminal polymers

used to prepare amphiphilic derivatives. (A) Branched poly-

(ethylene glycol)-COOH [35]: (B) poly(acryloy1 morpholine)-

COOH [36]; (C) poly(vinyl pyrrolidone)-COOH [37].

which was then polymerized in water with a free-radical initiator as described in [36]. To prepare PVP-COOH (Fig. 4C), PVP-OH was first obtained by chain transfer free-radical polymerization of N-vinylpyrrolidone in isopropox- yethanol. PVP-OH was then converted to PVP- COOH by activating the hydroxy group with 4-nitrophenyl chloroformate and subsequent coupling with glycine as described in [37]. The molecular weights of PAcM-COOH and PVP- COOH were ca. 6000. To make these polymers amphiphilic and capable of anchoring into the liposomal membrane, each carboxyl-terminated polymer was then derivatized with a single phosphatidyl ethanolamine (PE) residue via carboxy group activation with dicyclohexylcarbodiimide (DCC) and N-hydroxysuccinimide (HSI) (Fig.

5). Liposomes (155-180 nm) were prepared by

the detergent dialysis method from a mixture of egg lecithin and cholesterol with the addition, when necessary, of 3 or 7 mol% of the corresponding amphiphilic polymer. Diethylene triamine pentaacetic acid-stearylamine was used for subsequent liposome labeling with “‘In. Biodis- tribution and blood clearance experiments in CD-1 mice clearly demonstrated that amphiphilic


0

HO N

3

_~O+&-%

0 0

HSI

3 pLJ

f

polymer-COOH - pOlYmer-COO- N 0.P-,O-(CH,),-NH-CO -Polymer

DCC W o-

0

Polymer-PE conjugate

Fig. 5. Reaction scheme of the modification of COOH-terminal polymers with PE.

PE-containing derivatives of PAcM, PEG2 and PVP can also provide effective protection to liposomes in vivo similar to PEG-PE, which agrees well with our theoretical considerations and previous experiments [23,25-27,32,33]. The extent of protective activity for different polymers depends on the structure and quantity of protecting polymer on the liposome surface.

Thus, PEG-PE and PEG2-PE both sharply increase the residence time of liposomes in the circulation. However, PEG2-PE seems to be noticeably more efficient than PEG-PE when the content of each polymer in liposome is 3 mol%, whereas at 7 mol% their effect on liposome blood clearance is practically identical. Half- clearance times for PEG and PEG2 liposomes with 3 mol% content of protective polymer are ca. 80 and 140 min, respectively, and for PEG and PEG2 liposomes with 7 mol% content of protective polymer ca. 230 min for both preparations (whereas half-clearance time for “plain” liposomes of the same size is only about lo-15 min). The results obtained can be easily interpreted taking into account the structure of both polymers. In PEGZPE, a single PE residue carries two PEG-5000 chains, whereas in PEG- PE, each PE residue is conjugated with a single PEG-5000 chain. As a result, when 3 mol% of PEG2 are used, the actual amount of protective PEG chains over the liposome surface is twice as high as for PEG-PE, which results in better protection. A similar phenomenon was described in [35] where PEG2 was shown to be a better steric protector for proteins than PEG. When 7 mol% of protected polymer are used, even single-chained PEG-PE can provide complete coating of the liposome surface and maximum steric protection, so there is no more room for an

additional effect of extra PEG chains of PEG2- PE.

The blood clearance data for PAcM-PE and PVP-PE also demonstrate a strong protective effect on liposomes. Half-clearance times for liposomes with 3 mol% of PVP-PE and PAcM- PE are ca. 40 and ca. 90 min, respectively, whereas for liposomes with 7 mol% of PVP-PE and PAcM-PE they are ca. 130 and 170 min, respectively. It is important to point out that half-clearance times for liposomes modified with 3 or 7 mol% of PVP-PE prepared by chemical modification of PVP-COOH with PE, almost coincide with half-clearance times for liposomes modified with similar quantities of PVP-palmityl of the same molecular weight prepared by free- radical polymerization with chain termination obtained by us in [32,33]. It shows that if the energy of interaction between the hydrophobic anchor of the amphiphilic polymer is sufficient to keep the whole polymer molecule on the liposome surface [33], the type of this anchor is of minor importance, and liposome circulation time will be determined by properties of the protecting polymer.

Recently, our observations were confirmed by the authors of [38], who prepared liposomes containing 5 mol% of distearoyl-PE covalently linked to poly(2-methyl-2-oxazoline) or poly(2- ethyl-2-oxazoline) (see Fig. 3). These liposomes also exhibit extended blood circulation time and decreased uptake by the liver and spleen. A similar observation was made in [39], where phosphatidyl polyglycerols (see structure in Fig. 3) were shown to prolong the liposome circulation time in vivo. Until now the chemistry of PEG is still the most elaborated [40,41], which permits to perform numerous coupling reactions

1.50 V.P. Torchilin. VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (199-f) 141-155

of PEG to liposomes and particles. Moreover, methods are developed that permit to couple targeted devices (antibodies) to the distal ends of the grafted PEG molecules [42,43]. The preparations described advantageously combine longevity and targetability similarly to what was described by us earlier in the case of co-immobiliza- tion of PEG and antibody on the liposome surface [9].

Thus, hydrophilic and flexible synthetic polymers such as linear and branched PEG, PAA, PVP, PAcM [7,23,33], polyoxazolines [38], and polyglycerols [39] when made amphiphilic by modification at one terminus with long-chain fatty acyl or phospholipid residue, can incorpo- rate into the liposome surface and make the liposome long-circulating. The protection effects observed are determined by the conformational behavior of polymer molecules in the solution and depend on both the length of hydrophobic “anchor” and the length and structure of the hydrophilic polymer chain. The scale of these effects might be interpreted in terms of the balance between the energy of hydrophobic anchor interaction with the membrane core and the energy of polymer chain motion in the water solution. All this agrees well with our theoretical model [23,27].

4. Protective polymers on the surface of nanoparticles and latexes

In addition to liposomes, synthetic amphiphilic polymers have also been used for steric stabilization of nanoparticles with the surface of hydrophobic nature in order to prolong their circulation in the blood and to alter their biodistribution. The detailed description of polymer-modified nanoparticulates is given in this issue in reviews by Moghimi, and Stolnik, Illum and Davis, so we have no need to discuss all the related information. We would like, however, to present some data demonstrating the principal similarities underlying the behavior of sterically protecting polymers on the surface of nanoparticulate carriers and liposomes.

There are several principal types of amphipathic polymers used so far for the coating of

injectable nanoparticulate carriers. All of them consist of a hydrophobic moiety which easily adsorbs on the surface and performs an anchoring function, and a hydrophilic moiety which protrudes into the solution and effectively protects particulates from interaction with plasma proteins in the blood upon intravenous administration.

Polystyrene latex particles, commercially available in a number of sizes in the range of 0.05-l ,um, are the carriers of choice for surface modification studies. Polystyrene latex particles have a more narrow size distribution as compared to liposomes, and, hence, might be a more conveni- ent model to study the phenomena connected with polymer modification of surfaces. The mechanism of protection is essentially the same as for PEG-containing liposomes - the conformational “cloud” of the polymer flexible chain protects the particle hydrophobic core from contact with opsonizing proteins [23,27]. Modification of the liposome surface with an amphiphilic polymer usually requires the incorporation of the hydrophobic moiety of the polymer into the liposome phospholipid bilayer, which is impossible in the case of solid particle modification. In the case of particulates, surface modification can be performed by one of the following two methods: (1) physical adsorption of a polymer on a particle surface; (2) chemical grafting of polymer chains onto a particle.

One of the most commonly used is the series of linear or branched copolymers of polyethylene oxide and polypropylene oxide (Pluronic/Tet- romc . TM or Poloxamer/PoloxamineTM). There are numerous studies on the coating of different surfaces with these polymers, including their use for protection of latex particles from the uptake by the reticula-endothelial system upon i.v. injection [44]. We described another type of amphiphilic polymers which were developed specifically for liposome membrane incorporation, but may be used for steric protection of nanoparticles as well - hydrophilic linear polymers with a terminal lipid or fatty acyl group. PEG-phosphatidyl ethanolamine (PEG-PE) is a typical representative of such polymers (see Figs. 1 and 3 and Ref. [7]). The third important type of amphipathic polymers is an amphiphilic copoly-


mer in which in the water medium the hydrophobic block itself is able to .form a solid phase (particle), while the hydrophilic part remains as a surface-exposed “protective cloud”. The example of such a structure is a block-copolymer of PEG and a polylactide-glycolide (PEG-PLAGA) [45,46].

The adsorption of Pluronic/TetronicTM or Poloxamer/PoloxamineTM surfactants on the surface of polystyrene latex particles proceeds via the hydrophobic interaction with the polypropylene oxide fragment. Interestingly, the absorption of poloxamer-type copolymers takes place only on the solid particles with a clearly hydrophobic surface; no interaction is detected, for example, between such copolymers and the liposome surface [44]. There have been numerous studies on blood clearance and biodistribution of these particles upon their coating with surfactants, including the use of surfactants for particle protection from the uptake by the re- ticulo-endothelial system upon intravenous injection [47,48]. Upon i.v. administration, hydrophobic particles of sub-micron size are believed to be opsonized with macrophage-recognizable but not yet identified serum proteins [49]. Similar to liposomes, protecting the particle surface with hydrophilic flexible polymeric chains results in a substantial decrease in phagocytosis by liver macrophages and subsequent prolongation of circulation time. Porter et al. [50] have demonstrated that the absorption of the above copolymers leads not only to the decrease of particle uptake by resident macrophages in the liver but, after coating with some specific copolymers, can redirect the injected nanoparticles to other organs. For example, the coating of 60 nm polystyrene latex with poloxamer 407 results in increased particle accumulation in bone marrow. The same group has demonstrated that an analogous procedure also helps substantially to alter the biodistribution of subcutaneously injected nanospheres. Coating of 60 nm diameter polystyrene nanospheres with certain poloxamer/ poloxamine copolymers results in their increased accumulation in regional lymph nodes. The opti- mal length of the copolymer polyoxyethylene block has been found to be 5-15 oxyethylene units. Non-coated particles have a tendency to

stay at the injection site while particles coated with longer polyoxyethylene-containing copolymers are not retained in the nodes and eventual- ly appear in systemic circulation [22].

Hydrophilic linear polymers with a terminal lipid or fatty acyl group, such as PEG-PE or poly(acry1 amide)-palmityl (PAA-P) [7,32], have also been tried as potential steric protectors for non-liposomal carriers. To investigate the behavior of such polymers on the surface of nanoparticulates, we have used PolybeadsTM polystyrene latex particles with a diameter of 100 nm. PAA-P (MW 12000) and PEG-PE (MW 5000) have been used to coat the surface of latex particles. The incubation of nanospheres with the polymers in water results in polymer attachment to the surface, which can be confirmed by the measurements of the particle size before and after incubation with the polymer. The particle diameter increase upon polymer adsorption was found to be ca. 5 nm for PEG-PE and ca. 20 nm for PAA-P. The results of biodistribution studies in mice demonstrated that PEG-PE-coated particles, as one can expect, stayed in the circulation for a long time (t,,, = 4 h). However, unlike PAA-coated liposomes [32], PAA-coated nanospheres cleared from the blood as fast as non- coated particles with a similar pattern of liver accumulation. At the same time, spleen accumulation of PAA-coated particles was substantially reduced compared to non-coated ones. Different effects of the same amphiphilic polymer coating on the behavior of different particulates in vivo have already been described. Moghimi et al. [44] have shown that poloxamer 407 can protect latex nanospheres, but not the liposomes of similar size, from RES uptake. This difference in the in vivo behavior has been explained by the different orientation of poly(acry1 amide) chains on the particle surface compared with the surface of liposomes, which results in a different degree of carrier protection from absorbing plasma proteins.

In a recent review on long-circulating liposomes [15] Allen points out the way of polymer attachment to the surface as a major difference of liposomes with poloxamers or PEG-PE. With poloxamers, the hydrophobic region lies perpendicular to the acyl chain region of the bilayer,

152 V.P. Torchilin. VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-1.55

whereas for PEG-PE the hydrophobic anchor is parallel to the acyl chains. As a result, the thickness of coating is considerably less for poloxamers on liposomes than for poloxamers on nanoparticles [51,52]. For liposomes PEG-PE works better than poloxamer F-108 [52]. Still, we are talking about the differences in anchoring various protecting polymers onto different surfaces, whereas the behavior of hydrophilic moiety in the surrounding water solution should remain pretty much the same.

Another important type of amphipathic polymers includes copolymers in which in the water medium the hydrophobic block is able to form a solid phase (particle), while the hydrophilic part remains as a surface-exposed protective “cloud”. The protective effect of hydrophilic polymers was demonstrated in a system with chemical “anchoring” of PEG chains onto the surface of polymeric nanospheres. For this purpose block copolymers of poly( lactic-co-glycolic acid) (PLAGA) and PEG have been synthesized. Using the solvent evaporation method, PLAGA polymer was demonstrated to form slowly biodegradable spherical particles of sub-micron size [46]. Applying this method to the PLAGA- PEG copolymer, one can prepare particles with an insoluble (solid) PLAGA core and a water- soluble PEG shell covalently linked to the core. Several polymer preparations have been synthesized in which the molecular weight of the PEG block was 20 kDa, and it was connected with the PLAGA block with 1:9, 15 and 1:4 PLAGA:PEG w/w ratios (Trubetskoy, Langer and Torchilin, unpublished results). The nanospheres were labeled by incorporation of hydrophobic ’ ’ ‘In-diethylene teriamine pentaacetic acid stearylamine into the PLAGA core during the solvent evaporation procedure. Blood clearance and biodistribution experiments in BALB /c mice have demonstrated that the protective effect of PEG in this system depends on the content of its block (Fig. 6). Clearance and liver accumulation patterns reveal one basic feature of the preparations under discussion: the higher the content of PEG block is, the slower are the clearance and the better protection from the liver uptake. The splenic uptake also reflects the particle size-dependent filtering effect. As has

A

0 20 40 50 80 100

time, min

0

0 20 40 60 80 100

time, min

Fig. h. Blood clearance (A), and liver (B) accumulation of

“‘In-labeled PEG-PLAGA nanospheres with different PEG-

:PLAGA w/w ratios in mice (Trubetskoy, Langer and Tor-

chilin, unpublished).

already been demonstrated for liposomes, the vesicles with a diameter exceeding 250 nm can be non-specifically detained in spleen [53]. 1:5 and 1:4 PEG-PLAGA nanoparticles were 265 and 315 nm in size, respectively, so the elevated splenic uptake for these preparations can be explained by their passive retention in the spleen. Similar effects on longevity and biodistribution of microparticular drug carriers might be achieved by direct chemical attachment of

VP. Torchilin, VS. Trubetskoy I Advanced Drug Delivery Reviews 16 (1995) 141-155

protective polyethylene oxide chains onto the surface of preformed particles [54].

5. Conclusion

Thus, steric protection of liposomes and other nanoparticulate carriers can be achieved by grafting their surface with many hydrophilic and flexible polymers including linear and branched poly(ethylene glycol), poly(acry1 amide), poly- (vinyl pyrrolidone), poly(acryloy1 morpholine), and some others. The protective effect of these polymers on nanoparticulate carriers including liposomes may be interpreted in terms of statistical behavior of macromolecules in solution, when polymer flexibility plays a key role. According to the mechanism proposed, surface-grafted chains of flexible and hydrophilic polymers form dense conformational “clouds” over the particle surface and prevent other macromolecules from interaction with the surface even at low concentrations of protecting polymers. The model permits to answer various questions concerning spatial distribution and density of polymer molecules on the surface of nanoparticulate carriers. The principal behavior of protecting polymers on the surface of liposomes and particles is similar, though the architecture of polymer coats may differ. The scale of protective effect can be considered as the balance between the energy of hydrophobic anchor interaction with the surface and the energy of polymer chain free motion in solution. The whole approach opens wide oppor- tunities for controlling the in vivo behavior of various microparticular drug carriers.

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