Stealth® liposomes: from theory to product

advanced

drug delivery reviews ELSEVIER

Advanced

Drug Delivery

Reviews 24 (1997)

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Stealth@ liposomes: from theory to product’ Boris Ceh”, Mathias Winterhalterb,

Peter M. Frederik”, Joseph J. Vallnerd, Danilo D. Lasicev*

“Departmenr of Chemistry, University of Ljubjuna, Ljubjunu. hBiozentrum Basel, Basel, Swirzerlond ‘Limburg Universiry, Maastricht, The Netherlunds “Seyuus Pharmaceuriccds, Menlo Purk, CA, USA ‘Liposome Con.sulration.v,Newark, CA, USA

Slovenio

Abstract The development of an effective anti-cancer liposomal formulation - doxorubicin in sterically stabilized liposomes will be discussed. We shall argue that for many tumors the necessary condition for an effective anti-cancer activity of systemically administered liposomal doxorubicin formulation is the long circulation life of liposomes in blood and stable drug encapsulation. Theoretical basis for stabilization of liposomes in biological environments and for the stabilization of drug encapsulation will be shown. When a formulation with acceptable stability was obtained it was tested in pre-clinical models and simultaneously scaled-up and it entered into clinical studies. After successfully passing all these tests, doxorubicin in sterically stabilized liposomes (DoxilTM by Sequus Pharmaceuticals, Inc., Menlo Park, CA) was approved by Food and Drug Administration and is commercially available since late 1995. Keywords:

Doxorubicin; DoxilTM; Drugs; Liposomes; Polymer

Contents

I. Introduction ............................................................................................................................................................................ 2. Steric stabilization of liposomes ............................................................................................................................................... 3. Some experimental results and discussion ................................................................................................................................. 4. Protein adsorption ................................................................................................................................................................... 5. Encapsulation and retention of doxorubicin in liposomes ............................................................................................................ 6. Drug delivery .......................................................................................................................................................................... References ..................................................................................................................................................................................

1. Introduction Liposomes

were considered

a drug delivery

sys-

*Corresponding author, 75 12 Birkdale Dr., Newark, CA 94560, USA. ‘Due to their invisibility to the body’s defence system sterically stabilized liposomes are often called also Stealth@ liposomes (a registered trademark name of Sequus Pharmaceuticals, Inc., Menlo Park, California. 0169-409X/97/$32.00 Copyright PI1 SO 169-409X(96)00456-5

0

165 166 168 171 171 174 175

tern of choice for systemic applications of anticancer agents due to colloidal size, easily controllable surface and membrane properties, large carrying capacity and biocompatibility. However, it took more than 20 years before largely unforeseen problems of liposome instability in blood circulation upon systemic application and problems with efficient and stable encapsulation of agents were solved [ 1,2] and first products become commercially available [31.

1997 Elsevier Science B.V. All rights reserved

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Typically, intravenously administered liposomes are taken up by the cells of mononuclear phagocytic system in dozens of minutes. Even faster, they can loose the encapsulated cargo, especially drugs with a low value of water/octanol partition coefficient. Following early observations that cholesterol increases liposome stability [4,5] it is now believed that stability of non-sterically stabilized liposomes scales with stretching elastic modulus of the membrane [6]. A partial solution is that by using mechanically the most cohesive bilayers, distearoyl lecithin or sphingomyelin, containing at least one third of cholesterol, could help increase the circulation times up to few or perhaps 10 h. However, very long circulation times could not be achieved and such formulations have inherent colloidal stability problems due to a neutral charge. Also, the question of mechanically strongest bilayers is still open and the use of lipids containing arachidonyl or longer fatty acids or polymerized bilayers has not been tested yet for this application. In addition to mechanical stability we considered biological stability of liposomes as a problem of colloid stability. Theoretically, such systems can be stabilized electrostatically or sterically [7]. It was proposed that the rapid clearance of liposomes is due to the uptake by the cells of the body’s immune system as well as disintegration upon interactions with plasma lipoproteins [8]. Proteins of the immune system adsorb onto liposomes immediately after injection and this opsonization tags them for the subsequent macrophage uptake. The molecular origin of these interactions are mostly long range electrostatic, van der Waals and short range hydrophobic interactions of liposomes with macromolecules in the blood. While electrostatic and hydrophobic interactions can be minimized by using neutral lipids and mechanically very strong bilayers, van der Waals attraction is ubiquitous. However, it shows a power law decay with distance. Van der Waals attraction can be, therefore, reduced by coating the liposome surface with inert polymers which prevent the close approach of proteins. The introduction of the spacer also causes modification of the aqueous solvent and can cause the change in the prefactor (Hamaker constant, A). Longest blood circulation half lives of liposomes, t, ,*, in mice were achieved by sterically stabilized liposomes with the composition: ““PEG-DSPE (polyfollowing

Delivery

Reviews 24 (1997)

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ethylene glycol-distearoyl phosphatidylethanolamine) I hydrogenated phosphatidylcholinel soy cholesterol _5/5.5/40 mol/mol). It is likely that the clearance of these liposomes is governed by the dissociation of PEG-lipid from the membrane and therefore using stronger hydrophobic anchors or polymerized bilayers can yield even longer circulation times. The second problem of doxorubicin liposomes is rapid leakage of encapsulated drug. Typically, halflives for passively encapsulated drug range from minutes in EPC bilayers to a dozen hours in the case of mechanically strongest bilayers, such as DSPC/ Chol. Efficiency of loading can be greatly increased by using a pH gradient loading technique [9]. Protons, however, are known to leak through the membranes and the gradients dissipate and the drug leaks out with time constants around 15 h from mechanically the most cohesive bilayers. For liposomes which can circulate in patients up to a week this means that they would reach their target empty and alternative techniques have to be used. One example is chemical potential gradient loading coupled with precipitation of the drug in the liposome interior [lo]. In such a case chemical potential drives drug from the exterior into interior where it, after protonation, precipitates. According to the law of mass action drug is continuously pumped in up to the equilibration which can in principle be as low as dictated by the solubility product of the precipitated drug. If one uses ammonium sulfate gradient (inside of the vesicles) the exchange of the neutral species occurs: ammonia leaks out where it is diluted and the external doxorubicin is driven inside where it precipitates [lo-131. The stability of the drug encapsulation was measured by various assays and it was found that, according to the theoretical predictions [ 13,141, ammonium sulfate gradient loading gives rise to more stable drug retention [ 15,161.

2. Steric stabilization of liposomes In the simplest picture it is possible to understand the stability of sterically stabilized systems as a function of two controlling parameters: polymer chain grafting density (D) and their length (degree of

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polymerization, N) [17]. We shall briefly mention scaling theory which was found to describe experimental data satisfactorily and we can use it to estimate the surface concentration of polymer needed for stabilization. Corona of attached polymer causes repulsion because the entropy of chains is reduced upon approach (loss of configurational entropy) as well as excluded volume and osmotic repulsion. Depending on the size of the polymer (R) and distance between attachment/grafting points (D) polymer can have different conformations. At very low surface coverages (D > R) the polymer forms either pancake like structures or inverse droplet-like structures, depending if it forms an,adsorption or depletion layer on the surface. At D > R the so-called mushroom conformation is present while for D 4 R polymer chains start to interact, forcing their extension into the so-called brush conformation [17]. Scheme 1 shows these conformations schematically. Scaling laws which enable the calculation of repulsive pressure above the grafted surface were reviewed recently ]6,17-211. Below we shall show that experimental data are in nice agreement with theory. In contrast to the established treatment of polymers at interfaces, the original qualitative proposal of

STERICALLY *

STABILIZED

LIPOSOMES

Lipid Mayer

Polymer

POL BR

Scheme I. Various conformations of liposome polymers (Drawing courtesy of S. Hansen).

surface

grafted

167

reduced bilayer adsorption [8] was extended into a simple theoretical model of a ‘statistical cloud’ [22]. The protein adsorption onto liposome surface was modeled by a random flight simulation. The protein is assumed to exert Brownian motion until it hits the liposome surface where it adsorbs. The polymer is assumed to be a linear chain composed of monomers which can freely rotate. The presence of polymer prevents adsorption. Stabilization is achieved when computer simulation shows effective surface coating by polymer. The main predictions of this model are: ( 1) the maximum inhibitory effect is achieved when the polymer layer coats the surface what can be achieved by increasing either chain length or grafting density. (2) Protection is a linear function of both variables and (3) inhibitory effect increases with polymer flexibility. Despite that qualitatively most of the arguments make sense, this model has several theoretical shortcomings. Firstly, the model does not account for the driving force for protein adsorption. Van der Waals interaction with polymer layer is neglected. The thickness of the protective layer is not considered and van der Waals attraction to liposome across a thin polymer layer is not taken into account. Secondly, polymer conformation is modeled to be invariant of its density, what was experimentally and theoretically shown not to be case [17,20,23-251. Thirdly, the polymer density inside the surface layer is homogeneous causing an isotropic pressure. This implies, that, once a protein penetrates the polymer it will either leave or adsorb with equal probabilities. Protein adsorption will therefore occur, only at a slower rate and, fourthly, theoretical models and experimental results predict and show that repulsive pressure is not linearly proportional to N and D, but rather there is a power law dependence. Coating the surface does not prevent protein adsorption by itself. It is the change in polymer conformation which gives rise to repulsion! Additionally, this model also cannot explain higher repulsive pressures at higher grafting densities [25]. Below, we shall show that measured repulsive pressures correlate very well with the scaling theory predicted power law dependence. To check this, however, detailed studies of protein adsorption should be performed as a function of D and N. In other words, this model, therefore, represents an irrelevant and trivial exercise because it predicts an

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isotropic protein cloud. It is exactly the change in polymer conformation which causes the force against the incoming protein and there is no linear transformation between mushroom and brush regime [17,20,21,23-251. Again, it is not the coating itself, but the change of coating polymer conformation due to incoming protein which causes a repulsive pressure. Poly (ethylene) glycol (PEG) is by far the most widely used polymer to impart steric stabilization [6]. The generality of polymer stabilization of liposomes was confirmed when other polymers were discovered which rendered similar biological stability. Poly-( 2-methyl-2-oxazolidine and poly-( 2-ethyl2-oxazolidine) with II = 50 attached to DSPE and incorporated into liposome bilayers at 5 mol% showed similar stability profile as PEG [26]. Poly(acryl)amide and poly(viny1 pyrrolidone) also increased blood circulation times of liposomes [27]. The inferior results, as compared to PEG are probably due to a weak hydrophobic anchor which caused release of these molecules from liposomes. Aqueous solubility of PEG-lipids [28] may be the cause of their clearance because dissociation half-lives are around 15 h for distearoyl chains. As expected, aqueous solubility critically depends on hydrophobic anchors and follows a scaling relation CCN3’5 with respect to the PEG size [16]. Recently the cmc of *“‘PEG-DSPE was measured by turbidimetry and a value of 5.6 PM was determined [29]. With the development of monoclonal antibodies and long-circulating liposomes, their targeting to particular cells became feasible. In addition to classical techniques to link antibodies, lectins or other ligands to PE, coupling to the far end of PEG chain was developed. Inert terminal methoxy group is replaced with a reactive functional group suitable for conjugation after liposomes were prepared containing such reactive polymer-lipid [30,31].

3. Some experimental

results and discussion

The hypothesis of increased repulsive pressure above membrane and reduced adsorption on the blood circulation lifetimes was tested by measuring the repulsive pressure between membranes with and without incorporated polymer-bearing lipid by using the osmotic stress technique. Fig. 1 shows that

I

I

I

I

I

-20

20

60

100

I40

Distance from Bilayer Center

(A)

Fig. I. Electron density profile for 2: 1 stearoyl-oleoyl phosphatidyl choline: cholesterol bilayers without (top) and with (bottom) 4.5 mol% 2”““PEG-DSPE. Two unit cells are shown for each profile. The position of lipid and polymers can be obtained from the stick figure models. (from [70], with permission.

bilayers containing PEG lipid show much larger interbilayer spacings and even upon strong compression the bilayers are still 4 nm apart as compared to surface unmodified bilayers which show practical collapse to the hard wall repulsion [24,25]. The interbilayer repulsion calculated from the repulsive pressure of surface attached polymer in a mushroom configuration from P = (5/2) kTNID2/a (al(h/2))8’3

(1)

and for N=44, a=0.35 nm, D = 3.57 nm, is in nice agreement with experimental data [ 191. For distances >h, the repulsive pressure is zero. Theoretical extension of this simple law led to the parabolic decay instead of steep single-step decay at h> h,. The sensitivity of these results, however, cannot distinguish between parabolic or step decay indicating that the simple scaling approximation is rather good [24]. Practically identical results (Fig. 2) were obtained also by the surface force apparatus [25]. These results also show increased repulsion with increasing amount of PEG polymer. Surface force measure-

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20 QO 18-

gQFb0

0

10

20

30 40 50 Bilayer Separation

60 (A)

70

80

90

Fig. 2. Repulsive pressure (force-distance profile) above bilayer with (circles) and without (squares) 4.570 ‘“““PEG-DSPE. Very good agreement between between surface force apparatus mesuremen& (open) and osmotic stree method (solid) were observed. (adapted from [24,25] and [70]).

ments have found reversible repulsive force at all separations and the thickness of the steric barrier was found to be controlled by the amount of added PEG-lipid [25]. Scaling analysis fits, which also describe the measured dependence rather satisfactory, were found to be for the mushroom regime F<,(h)IR = 1.6 (2 kTID2) [(R,lh)“‘3

-

l]

(2) 1

where the numerical prefactor is close to expected unity [25]. At higher grafting densities, i.e. in the brush regime, the force could be described by F<,(h)IR = (16 kT r h<)/ (350”) + 5 (h/2h,.)7’4

-

121

[7(2h, /h)5’4 (3)

.I 0

0.1

3

I

4

_

I

5

6

7

8

9

Distance, D (nm) Fig. 3. Theoretical fits for mushroom and brush regime. Simple scaling analysis (dG curves) accurately describe the repulsive force as a function of distance (adapted from [25]).

where h, = D (R,lD)5’3

(4)

and Fig. 3 shows a comparison of experimental points with a mushroom and brush expression of force-distance profiles. Force between two cylindrical surfaces (F,) and repulsive pressure (P), as measured by the osmotic stress technique, can be calculated using Derjaguin approximation F<(h)lR = 2~

i

The agreement

P (h) dh with theory

(5) was good up to the

distance of two polymer layers. At larger separations, the theory, however, predicts a steep decrease which was not observed. Obviously, by simply taking into account polydispersity of the polymer as well as of polymer configurations would alleviate this inconsistency. Good agreement with this simple scaling concept was observed also in the regime of interacting mushrooms [32]. We can use the above scaling concepts to estimate the thickness of the polymer layer needed for effective stabilization. In principle it can be regulated by either surface grafting density and polymer chain

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length. In practice, however, shorter PEG chains attached to lipids are preferable due to lower aqueous solubility of these lipids and therefore higher retention in the bilayers. The thickness of the coating can be expressed in a good solvent as h, = DN@zID)“‘~ where N is degree of polymerization, a size of a monomer and D the mean distance between the grafting points. Using scaling arguments we can estimate polymer thickness for effective stabilization: a general criterion becomes then A (b/h,) lo-18% cause liposome solubilization and formation of mixed micelles [35-381, and that pure PEG-lipids form small micelles 1391, lipid monolayers of pure PEG-lipids were shown to be stable up to >45 mN/m at area of ca 0.75 nm2 per molecule [34,40-421. Monolayer studies may be a little tricky, because free PEG is

Reviews 24 (1997) 165-I 77

also a surface active agent. Its collapse pressure is around 8-15 mN/m (increasing with increasing N). This may mean that a sufficient pressure from the lipid is needed to push PEG into the subphase while at lower pressures there may be coexistence of lipid islets and surface spread PEG [41,42]. Furthermore, spherical vesicles prepared from pure PEG-lipids were observed by EM [M. Kasbauer, University Wurzburg, Physics, Diploma thesis, unpublished]. Obviously, more work is needed to reconcile these observations. It is possible that PEG-lipid undergoes phase transformation from random coil into helix, as it happens in the solubility gap in the phase diagram of PEG in aqueous solutions. This transition may be exploited for changing very stable liposomes into fusogenic upon triggered conformational transition of PEG chains. Indeed, it was reported that PEG-lipid liposomes induced fusion with mammalian cell membrane. Egg yolk lecithin liposomes containing 30-60 mol% of PEG-lipid (2-[-w-hydroxypoly(ethylene)-a-yl] - I,3 - bis-(dodecyloxy)propane), with degree of polymerization N = 15, and in their interior plasmid pActDTA, which is fatal to HeLa cells. Among many controls only PEGylated liposomes reduced the number of cells indicating that plasmid entered cells by a process different from endocytosis which would degrade the fatal plasmid [43]. Experimental measurements using osmotic stress technique, and surface force apparatus have shown that polymer extends, in agreement with theoretical calculation, about 5 nm above the bilayer. Several other groups have recently reported similar results by using electrophoretic mobility measurements. Hydrodynamic thickness of the polymer coating can be estimated from the zeta potential which, at the shearsurface equals electrostatic potential. The electrostatic potential profile is obtained by the Poisson Boltzmann equation and can be calibrated by using PG liposomes [44]. Such measurements showed that PEG coat extends about 4 nm above the bilayers at 5 mol% of zcrtr”PEG on the liposome surface [45]. Molecular weight dependence [44] showed reasonable agreement with expected power law behavior (v = around 0.6). A thorough analysis of the polymer thickness as a function of grafting density was done by Hirota and coworkers. They used PEG-dimyristoy1 glycerol with N = 44. By plotting the logarithm of zeta potential vs. Debye length they could evalu-

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ate the thickness of the polymer coating. The result of 3.2 nm for 5 mol% of PEG-lipid is in a reasonable agreement with other data and theoretically expected 3.5 nm [46]. Similar study was performed also using a series of PEGylated cholesterols. PEG layer thickness of N = 15 PEG-cholesterol was around 3 nm for 10 mol%. Thickness was found to decrease with increasing ionic strength indicating that water activity decreases with increasing ionic strength. In PEGlipids, however, such behavior was not observed ]471. PEG coated liposomes were shown to be stable against free PEG induced fusion [48]. Liposomes containing few mol% PEGylated lipids are stable against free PEG (30 wt% ““PEG) induced fusion while liposome containing DSPE instead PEG-DSPE were fused. Also the leakage of encapsulated dye was inversely proportional to the fraction of PEGylated lipid in the membrane.

4. Protein adsorption Protein adsorption on lipid membranes is basically driven by long range forces, such as electrostatic and van der Waals attraction as well as short range hydrophobic interaction. Although this subject is widely investigated neither a qualitative nor any quantitative understanding was achieved to date. The main problem is the lack of quantitative values of interaction constants, such as the Hamaker constant. Even more complex is the case of surface associated polymer. A qualitative theory was introduced few years ago which took into account basic forces [49]. Despite the elegant approach, however, due to the lack of experimental values for the hydrophobic interaction and Hamaker constant, no quantitative predictions could be made. Even less was done experimentally in the case of protein adsorption on sterically stabilized liposomes. While many studies have indirectly shown that PEGylated liposomes are less amenable to protein adsorption, and that the amount of adsorbed protein is inversely proportional to blood circulation times of such liposomes with various 1501, interaction proteins has not been yet systematically measured. Nor were corresponding monolayer studies performed. In black lipid membranes it was shown that insertion of PEGylated lipids reduced interactions of

lipoproteins with the lipid bilayer proportionally to the polymer size and grafting density. Interaction of lipid membranes with proteins was measured indirectly via the kinetics of pore formation during the irreversible electric breakdown of the membrane [51]. Around 1 mm2 large black lipid membranes, with and without PEG lipids were formed. The rupture velocity of the film was measured upon addition of lipoprotein and it was shown that the presence of 5 mol% of PEG-lipid inhibits lipoprotein adsorption on the membrane. In planar lipid membranes (also called black lipid membranes) it was shown that insertion of as little as 1% of 2’10”PEG-DSPE into the lipid solution reduced interaction of lipoproteins with lipid bilayer. More recently (Winterhalter et al., in preparation) it was shown that the addition of 4 mol% of the same polymer-lipid reduces the streptavidin binding to biotinylated lipids. This interaction is one of the strongest known and has binding constants of 10” M-‘. In this investigation the interaction of lipid membranes with proteins was measured indirectly via their influence on the kinetics of pore formation during irreversible electric breakdown of membranes. The rupture velocity was shown to depend on the lipid composition and is altered in the presence of membrane bound proteins and surface grafted polymers (it scales with the surface mass).

5. Encapsulation in liposomes

and retention

of doxorubicin

In addition to the increased stability of liposomes in blood circulation, the second necessary condition for an effective tumor targeting drug delivery vehicle is stable drug encapsulation. Because pH gradient loading is compromised due to proton leakage we used the exchange gradient technique coupled with precipitation of protonated drug with pre-encapsulated anion. Herein we shall very briefly review the underlying theory of this process. We shall briefly present the model, detine conditions and show the results. For details, however, the reader is referred to [ 13,141. The interior of liposomes contains a solution which is a mixture of Bronsted acid-base species. Before loading, the outer solution is exchanged for another solution that, in general, also contains an optional number of

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Bronsted acids, bases and salts. After this exchange a number of concentration gradients is established. Chemical potential differences which establish across the bilayer drive the corresponding neutral molecules, which can permeate through the liposome membrane, into the opposite compartment until their concentrations (chemical potentials) on both sides of the membrane are equal. The permeation of all charged species is assumed to be negligible. Assuming that the equilibrium acid-base constants in the inner and outer solution are the same, and taking into account also a precipitate formation inside the vesicles, three nonlinear x, y and R” (proton concentrations inside and outside of liposomes, and solubility product, respectively) dependent equations can be obtained, which describe the state in the liposome interior, exterior and the state of the precipitated drug. They take into account all possible initial chemical parameters in both compartments (outer/inner volume ratio, starting concentration of acids/bases/salts/drugs, their acidity constants and chemical composition as well as solubility product of the precipitated form of the drug). The master equation for the inner solution is [ 131:

2

k#k*

gkO(Y) 'k0

0

Akku

Ii-

I

+.&o(Y)

.A&>

f

c!5no(Y) ‘n, ’ .t&) + Kf,,(Y)

Finally, as:

the solubility

product

L, can be expressed

“i,

1

L,,= CL0 U&Xl

+ RI,,) +

(;)&Cyi (8)

Here, ckO and c,,, are the starting-point concentrations of species k in the outer solution and m in the inner solution7 .& f,,, &f;nl? f,o> gk@ gkl? &,1, &o are concentration and charge functions of the corresponding species k and m, respectively, A, and B,, the factors which ‘allow’ (= 1) permeation or leakage of neutral species kl or ml or not (=O). Three parameter, x (inner H+-concentration), y (outer H+concentration and R II,*, (concentration-ratio of membrane-bound and solution-remaining sister species kj*) can be obtained from above non-linear equations by a numerical Newton-Raphson method. Furthermore, loading/unloading efficacy of each species released into the outer compartment, pz, and loaded into the vesicles, az4, can be expressed as follows:

(9) g,,,(Y)

+

ck*O k,

+ U,(Y)

0

=

;

1k*-‘.M4U

+Rfl,d

+

./L,(Y) ‘4 ak*

0

fk*I(x)(’

and for the outer solution

one obtains:

s,,w c Crnl m+nl* J;,,(x)+ B,,,K,,(y.Lo(.v) g,, +

c k#k*

(4

AkckO k&A4

+

(f>im-

g,*,(d

+

‘k*O b=i&)(

1 +

R:,

*I

Ifko(y)

+ R&J

.fi&)(l = +

Rtj*,)

+

K,(:)“*Jf;*O(.Y) (10)

Solving the above equations for doxorubicin one gets the encapsulation ratio. These are general equations which can be simplified for some concrete examples. The Master equations describing a concrete case of the loading of doxorubicin (DOX, added in the form of doxorubicin hydrochloride) into the liposomes, for example, have the following forms; for the inner solution [14]:

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et al. I Advunced

Drug Delivery Reviews 24 (1997) 165-177

K,,K,2/.r2 X

173

Il+K,,/rtKL,/lZ)

K,~'~ItK,,i+K,,K~,ir')(ltR')t ~'K~'(,tK,,l~+K,,K,1/~i)(ltR'.)+($)(l+~,,l.~tK,,K,li~')

X(&l - I (”

the percentage of the encapsulated calculated from

2K,,/.r

I +K,,/x I

+

t2c, (ItK,,/r)+K,

n = (I + K,,/r t K,,K,,Lr’NIt /?“It K,(t)lI t

can be

K,,/s

+ K,,K,,/,?)

(14) I - K,,K,J,v2

Cl (I tR”)(I

DOX

(11)

for the outer one:

K,’

(13)

tK,,/x+K,,K,21X’)t(1 +K,,lytK,,K,Jy’)

(12) and for the solubihty product formed inside the liposomes:

of the precipitate

4, = Cl K,.‘(ItK,,/.~tK,,/~~~(ltR”~t 1 (ItK,,i+K,,/:‘) i.J

Here, c, is the initial concentration of the doxorubicin, C, =O, c3 is the starting concentration of the ammonium sulfate in the liposomes and a = 0. The corresponding &,‘s are: for DOX p,, =8.2, 17,~= 10.2; for ammonia, p7, =9.26, and for sulphuric pd2= 1.92. The solubility product for acid, (DOXH),SO, was measured and estimated to be 1.1 X lo-‘. Fig. 4 is an electron micrograph which clearly shows the presence of precipitated agent in the liposomes. Fig. 5 shows theoretical tit and experimentally determined points of doxorubicin loading from Eq. (14). Solving these equations for x and _y yields the pH values inside and outside of the liposomes and allows calculation of loading efficiency for various experimental conditions.

Fig. 4. Cryoelectron micrograph of doxorubicin precipitated in sterically stabilized vesicles. The presence of gel-like bunches of fibers, separated by about 2.7 nm) can be observed. The same sample gave small angle X-ray diffraction pattern with periodicity between fibers of 2.8 nm [ 101. This indicates that the enclosed drug is precipitated in the form of fibers ([doxorubicin’],-sulfate=),, which are hexagonally packed into bundles with interfiber separation of 2.8 nm.

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7.5 .

*

7.0 .. ; 6.5 .~’ P ti 6.0 -$5.5

-~

e 5 5.0 -4.5 1 0

CLntratlon

00X.&l

IO of

Fig. 5. Theoretical

fits and experimental

4.0 1 0

20 (mm)

points of doxorubicin

6. Drug delivery The first and still by far the largest application of sterically stabilized liposomes is in anti-cancer chemotherapy. The major problem in normal chemotherapy are toxic side effects of the treatment with very potent drugs. Normally, only less than about 1% of the intravenously administered dose of the free drug reaches the target cells. The remaining majority causes toxicity in various organs. The early idea of using liposomes was based mostly on their reduced toxicity. Indeed, the same drug doses were found to be several times less toxic when encapsulated in liposomes. Many researchers, however, overlooked the simple fact that the reduced toxicity was simply a consequence of the quick uptake of liposomes by the cells of the immune system, located mostly in the liver and spleen. Reduced toxicity was, therefore, accompanied by reduced efficacy of the treatment and with a possible exception of few very specific cases, such treatments resulted in diminished therapeutic efficacy [ 11. Small, sterically stabilized vesicles, however, were shown to extravasate into tumors [24,52] and the amount of extravasated liposomes was shown to be proportional to the blood circulation time [24,53]. The effect can be explained by the increased permeability (and often mechanical defects) in tumors which allow extravasation and, on the other side, damaged lymph drainage from the site and this cannot clear the extravasated particles. After many very encouraging preclinical data [5457, reviewed in 581 and safety studies the formulation entered clinical trials. In parallel the manufacturing process was scaled up and all the quality controls

loading

5 10 15 20 Concentmtionof DOXHCI (mm)

into pre-formed

liposomes.

(From ref. 14, with permission.)

and analyses defined and standardized. Scheme 2 shows the manufacturing process and some details are explained in the caption. After the anti-cancer activity in Kaposi sarcoma and safety of the formulation was proven in clinical trials [59-611 this drug was approved for sale and is on the market since late 1995. Currently the same formulation is studied in other tumors and other anti-cancer agents in sterically stabilized liposomes are being developed. This closes the chapter on the development of this type of formulation. Coupling polymers with liposomes, however, can achieve other beneficial effects and the future research may concentrate on those. Recent developments involve ligand targeting [ 12,62-651 (providing that problems with immunogenicity and site accessibility are solved), postliposomal steric stabilization by insertion of PEGlipid from micelles [29] and lamellar - hexagonal phase transition induced by thyolysis of PEG chain from the DOPE anchor [ [66], see below]. In addition to stabilization and targeting, surface active groups can induce vesicle aggregation, fusion among themselves or with cells and control leakage and possibly phase behavior. Inclusion of lipids with time dependent conjugated charge or polymer can add further ways of regulating colloidal and mechanical liposome stability and phase behavior. For instance, bilayers composed from micelle and inverse micelle forming structures, such as PEG-lipids and dioleoyl phosphatidylethanolamine, form stable bilayers. However, if PEG polymer is cleaved (by hydrolysis of a bond with known reaction kinetics, by photo- or pH-induced trigger) or by simply the dissociation of the whole molecule from the bilayer (due to shorter acyl chain(s), for instance), phase

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phase transition) can be induced also by free polymers which change their conformation as a function of pH [68,69]. We believe that next progress in liposome applications will be based on such developments, which couple properties of two different systems into a synergistic assembly.

References [I] Lasic, D.D. and Papahadjopoulos, Revisited. Science 217, 1245-1246.

D. (1995)

Liposomes

[2] Lasic, D.D. (1993) Liposomes: from Physics to Applications. Elsevier, Amsterdam, pp. l-575. [3] Lasic, D.D. (1996) Doxorubicin liposomes. Nature 380, 561-562.

in sterically

stabilized

[4] Papahadjopoulos, D., Nir, S. and Okhi, S. (1972) Permeability properties of phospholipid membranes: effect of cholesterol and temperature. Biochim. Biophys. Acta 163, 240249. [5] Gregoriadis, G. and Senior, J. (1980) The phosphoilipid component of SUV controls the rate of clearance of entrapped solution from the circulation. FEBS Lett. 119, 4341. [6] Lasic, D.D. and Needham, D. (1995) Stealth liposomes prototypical biomaterial. Chem. Rev. 95. 1601-1628.

Scheme 2. Schematic presentation of liposome manufacturing. In the case of the DoxilTM production lipids are mixed and dissolved in FDA and safety committee approved organic solvent which is injected into an aqueous solution of ammonium sulfate. Upon hydration the organic solvent is dialfiltered away and large multilamellar vesicles are sized down, preferably by an extrusion technique. After the exchange of external solution empty liposomes are loaded with doxorubicin which is added as a concentrated solution into the system. Loading is so efficient that free drug removal is not necessary. Some buffers may be added and the product is than prefiltered. sterile filtered, filled into vials, sealed and labelled. For convenience of medical persona1 and due to the robustness of physical and chemical stability and strong drug retention, the product can be supplied in liquid form which is the preferred form in parenteral medical practice.

transition occurs and encapsulated material is released [67]. In addition to lamellar hexagonal phase transition, one can induce lamellar - micellar by cleaving acyl chains of the lipids. By using labile PEG linkage one can induce steric destabilization, i.e. transition stealth regular liposome. Fusion, leakage or liposome dissolution (lamellar - micellar

a

[7] Napper. D.H. (1983) Polymer Stabilization of Colloidal Dispesions, Academic Press, New York. [S] Lasic, D.D., Gabizon, A., Huang, K., Martin, F.J. and Papahadjopoulos, D. (1991) Sterically stabilized liposomes: a hypothesis on the molecular origin of extended circulation times. Biochim. Biophys. Acta 1070, 187-192. [9] Nichols, J.W. and Deamer, D.W. (1976) Cathecolamine uptake and concentration by liposomes maintaining pH gradients. Biochim. Biophys. Acta 53, 37-46. [IO] Lasic, D.D., Frederik, P.M., Stuart, M.C., Barenholz, Y. and Macintosh, T.J. (1992) Gelation of liposome interior. FEBS Len. 312, 255-258. [I l] Haran, G., Cohen, R., Bar, L.K. and Barenholz, Y. ( 1993) Biochim. Biophys. Acta 1151, 201. [I21 Bolotin, E.M., Cohen, R., Bar, L., Emanuel, N., Ninio, S., Lasic. D.D. and Barenholz, Y. ( 1994) Ammonium sulfate gradients for efficient and stable remote loading of amphipathic weak bases into liposomes and ligandoliposomes. J. Lip. Res. 4, 4555479. [I31 Ceh, B. and Lasic D.D. (1995) A rigorous loading. Langmuir I I, 3356-3364.

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