COMPARISON of "STERICALLY STABILIZED" LIPOSOME TECHNOLOGY
and LCM TECHNOLOGY (FILMIX®) for ANTI-CANCER DRUG DELIVERY
The preparation of liposomes, with entrapped solutes, was demonstrated for the first time in 1965 in a published paper (J. Mol. Biol. 13:238-252, 1965) by Prof. A.D. Bangham of the United Kingdom. In early 1995, scientists met at a conference held in Cambridge,U.K. to celebrate 30 years of liposome research [1,2].
As B.F. Haumann describes in a review [2] of the field: “Liposomes are microscopic spheres made from fatty materials, predominantly phospholipids. Because of their similarity to phospholipid domains of cell membranes and an ability to carry substances, liposomes can be used to protect active ingredients and to provide time-release properties in medical treatment.
“Liposomes are made of molecules with hydrophilic and hydrophobic ends that form hollow spheres. They can encapsulate water-soluble ingredients in their inner water space, and oil-soluble ingredients in their phospholipid membranes. Liposomes are made up of one or more concentric lipid bilayers, and range in size from 50 nanometers to several micrometers in diameter” [2].
As explained in B.F. Haumann’s review [2], “The idea of using liposomes for drug delivery has been around since the early 1970s. The goals are to protect the body from unwanted side effects of various drugs and, when made to be targeted to specific tissue, to achieve desired concentrations of these drugs at a target site.”
“... Three companies established during 1980-1985 are among the leaders [in the United States] in the field of medical applications for liposomes. These are Liposome Technology Inc. (LTI) in Menlo Park, California; The Liposome Company Inc. in Princeton, New Jersey; and NeXstar, based in San Dimas, California. Formed somewhat later, Argus Pharmaceuticals in The Woodlands, Texas, also is a major player” [2].
Despite much research, “one hurdle has been to find ways to prevent the body from breaking down liposomes while they are still in the bloodstream and before they reach a tumor site. Conventional liposomes are limited in effectiveness because of their rapid uptake by macrophage cells of the immune system, predominantly in the liver and spleen. ‘They wouldn’t get to where you want them to go,’ according to an LTI spokesman” [2].
With regard to the short in vivo half-life of conventional liposomes, “Researchers at a number of companies have overcome this obstacle by designing liposomes that are nonreactive (sterically stabilized) or polymorphic (cationic or fusogenic). LTI, for instance, created its 'Stealth'® [i.e., sterically stabilized] fatty acid, a phospholipid with polyethylene glycol (PEG) attached to prevent the liposomes from sticking to each other and to blood cells or vascular walls. ‘Stealth’ liposomes appear to be invisible to the immune system and have shown encouraging results in cancer therapy" [2].
“Thus, coating liposomal vesicles with a hydrophilic polymer such as PEG reduces uptake by the liver. As a result, coated liposomes remain in circulation longer than conventional liposomes. Also, by incorporating targeting ligands on the surface of the liposomes, it is possible to direct them to certain organs” [2].
“Liposomes have had a rocky history, however, according to the LTI spokesman. ‘It is important to point out that the technology had its 30th birthday before a single product has been approved in the United States,’ he said.” In addition, “The first liposomal drug to gain approval anywhere was ... in 1990 in Ireland.” As of mid-1995, “Japan, meanwhile, has yet to approve a liposomal drug product” [2].
“Disappointments in performance include lack of a general strategy for tissue-targeting, drug concentrations which are sometimes lower than desired, difficulty of large-scale production, and unpredictable shelf life. ... A big hurdle is that encapsulating drugs in liposomes ‘currently is not a one-size-fits-all’ operation, according to LTI. ‘Instead, loading involves custom building the liposomes.’ LTI’s experience in creating its DOX-SL product ‘has humbled us in the difficulties of building a liposome to fit another drug,’ the LTI spokesman said” [2].
Additionally, in some applications, less expensive starting materials, such as polymers, are being used instead of phospholipids. ‘We have looked at many novel phospholipid-based drugs but, because certain phospholipids can cost $100 to $150 a gram, they are impossible to use for some applications. When we did the math, we could not rationalize their use,’ T.J. Pelura said [senior director for Alliance Pharmaceutical Corp. in San Diego,California]” [2].
Moreover, “IGI Inc. does not use phospholipids because they are very expensive, inherently unstable, very difficult to scale up to commercial production, and difficult to engineer. ‘For instance, you have to find a drug that can be used with them for certain niche applications,’ D. O’Donnell [IGI’s vice pres. of medical affairs] said, pointing out, ‘All of the liposome companies are concentrating on high-end chemotherapy agents’ ” [2].
Continuing modifications to the PEG coupling method, used in the production of stealth liposomes, are being employed in an effort to improve the reproducibility and usefulness of this method [3]. “This avoids both waste of PEG-lipid and loss of interior space. The latter can be substantial where high molecular weight contents are to be carried (with entrapped proteins an estimated 68% or 34% of interior space of a 100 nm liposome would be lost with PEG 5000 and PEG 1900, respectively). Allen et al. [4] have also reported foaming and micelle formation when attempting to incorporate PEG-phosphatidylethanolamine (PEG-PE) into liposomes with a lower incorporation level than PEG-PE added: with PEG-1900-PE at 10 mol% only 5.7, 5.0, 6.8, 6.5 mol% became incorporated into liposomes and the comparable figures for PEG-5000-PE were 6.9 and 7.3 mol%. This is an undesirable source of variation in PEG-liposome formulations” [3]. Moreover, “other methods leave part of the activating group between the PEG and the lipid (a coupling moiety). Even small coupling groups can affect functionality. In PEG-proteins, such coupling moieties have introduced between one to four problems being either: antigenic; biochemically and enzymatically degradable; inherently toxic; or changing the net charge of the target molecule” [3].
In contrast, the production process for research-grade LCM (i.e., lipid-coated microbubbles) does not require the use of PEG-lipids or any other synthetic polymers for steric stabilization (cf. above and sect. III), does not involve the use of phospholipids of any type (cf. sect. IV), and does not even require the use of sonication. Instead, the monolayer-coated LCM are produced simply by mechanical shaking of an aqueous suspension of (mostly) naturally occurring, nonionic lipids. The monolayer-forming lipid mixture used to manufacture the research-grade LCM is comprised of simple, “off-the-shelf” saturated glycerides and cholesterol esters, in a fixed ratio and of specific chain lengths [5]. The same nonionic lipids, used in the same ratio of lipid species, has been successfully employed when (non-covalently) incorporating either different fluorescent dyes (i.e., DiO [6] or Filipin [7]) or an anti-cancer drug (Taxol [8]) into the LCM monolayer structure. This degree of versatility of (hydrophobic) agent delivery, for one fixed mixture of simple “off-the-shelf” lipid ingredients, by LCM is quite different from the picture quoted earlier for liposomes (cf. sect. IV). Moreover, since formation of the monolayer-stabilized LCM is solely by molecular self-assembly [9], no organochemical reactions or derivatization is needed to assemble the LCM structure -- which avoids all the potential problems and complications associated with PEGylation of liposomes (see above). Finally, once the LCM are self-assembled, purification of these research-grade LCM during manufacturing consists of only a single filtration step [10]. The LCM manufactured by this simple method have been employed successfully in all Filmix® animal-efficacy studies published to date (see citations below).
(NOTE: A more detailed comparison of the manufacturing steps involved in the above two separate research-grade products, i.e., stealth liposomes versus Filmix®, can be made available to a potential licensee of Filmix® once license negotiations have reached an advanced stage.)
LCM have similarities with, and have been modeled after, the naturally occurring and very stable microbubble populations found in oceans and other natural waters. (Natural microbubbles represent an example of both self-assembly and nanoscience in nature, i.e., self-assembling nanoparticles that are widespread in natural waters.) Various physicochemical and biochemical studies on these long-lived natural microbubbles were conducted, and published, in the 1970s and early 1980s [11-19]. This university-based research contributed to the founding of the independent company Cavitation-Control Technology (Cav-Con) in 1979.
These early studies revealed that while different classes of organic matter coated and thereby stabilized natural microbubbles, it was actually an underlying monolayer of lipids (surrounding the gas bubble) which provides the long-term stability to such coated microbubbles [9]. The key structural categories of these stabilizing natural lipids were eventually identified, and then “off-the-shelf” substitute lipids were used instead to produce stable artificial microbubbles. All of the artificial-microbubble-formulation work was conducted, and initially patented [5], in the late 1980s. These artificial lipid-coated microbubbles were found to be very long-lived, lasting over 6 months in vitro [20]. In view of such longevity and the fact that the preferred formulation was comprised of simple saturated glycerides and cholesterol esters, these artificial lipid-coated microbubbles (LCM) then became a good candidate for potential medical applications.
The medical applications of LCM were explored and published during the 1990s [6,8,21-26], and such Filmix®-based safety and efficacy studies continue presently. An encouraging early finding was that after intravenous injection into animals, these LCM easily traverse the pulmonary microcirculation. In addition, unlike most other microbubble preparations reported in the literature then (and many even now), the LCM apparently survive literally many dozens of passes through the heart [22-26]. Furthermore, much evidence was obtained from numerous studies [6,8,22-26] that LCM are pliable enough and sufficiently small to pass across the known fenestrated capillary walls of tumor-tissue microcirculation [27-29] (see below).
As concerns these two methods for drug delivery to tumors, Barbarese et al. [6] recently compared the following research findings: “The longevity of LCM in the circulation is ... [a property] shared by small unilamellar liposomes (0.2 um diam.) [30]. Conventional liposomes administered by i.v. injection are usually cleared up by the reticuloendothelial system. Indeed, circulating macrophages and resident macrophages of various tissues rapidly phagocytose liposomes leaving only a few to reach their target [31]. However. a new class of liposomes, sterically stabilized liposomes, also referred to as stealth liposomes, has emerged which exhibit reduced uptake by the immune system [32]. When sterically stabilized liposomes were used to deliver encapsulated doxorubicin to tumors, it was shown that the maximum tissue accumulation of the drug occurs at 24 hours [33]. In contrast, our data suggest LCM to be rapidly removed from the circulation by tumor cells. The maximum accumulation of LCM in the tumor area occurs within the first 30 min after administration, and LCM can no longer be detected in the same region or in any brain region after 6 hours” [6].
“There is an increased blood flow in brain tumors [34], and the blood-brain barrier is leaky in and around 9L tumors because the blood vessels associated with the tumors are fenestrated [35]. This would allow extravasation of LCM. Similarly, extravasation in the extracellular space within tumors of sterically stabilized liposomes carrying doxorubicin occurs and is believed to be the chief reason for their efficacy [36]. Once in the tumor area, LCM may remain there because of an affinity for tumor cell-surface components" [6]. These components are more abundant on tumor cells than on normal cells. The mechanism for the uptake specificity appears to involve the LDL (B/E) receptor, and multi-ligand lipoprotein receptor, endocytic pathways [36a], -- and "at least 4 different types of tumor cells (C6 glioma, 9L gliosarcoma, Novikoff hepatoma, and Walker 256 carcinosarcoma) [25,26] do interact with LCM in a preferential manner suggesting that LCM affinity may be for tumor cells in general” [6]. This belief is further strengthened by separate studies in dogs [37] in which it was observed that “When injected intravenously, the LCM are effective in targeting ... SPONTANEOUS malignant tumors of the prostate, liver, and spleen in dogs [37]” [25].
Returning now to the study of Barbarese et al. [6], they further report: “Most 9L and C6 tumor cells appear to have LCM in their cytoplasm 2 min after LCM administration, while only a small proportion of sterically stabilized liposomes appears to be present intracellularly even after 24 hr of administration [33,38].” This observation of a markedly greater rate and extent of accumulation of LCM, in comparison to stealth liposomes, within the cytoplasm of tumor cells in vivo draws attention to another remaining hurdle faced by liposome researchers today. As pointed out recently by the Centre for Drug Delivery Research, School of Pharmacy, University of London: “A major challenge in terms of controlling liposome behavior in vivo is to persuade vesicles to deliver their contents selectively into the cytoplasm of cells through fusion” [1].
The LCM internalization process, carried out comparatively rapidly and efficiently by various types of tumor cells in vivo, has also been studied from a cytostructural standpoint both in vivo and in vitro by Barbarese et al. [6]. They report: “Once in contact with tumor cells, LCM are internalized bringing their content into the cell cytoplasm. This was observed by confocal laser scanning microscopy both in the brain in vivo and in cells in culture. Serial optical sectioning of cells that have been incubated with DiO-LCM shows that fluorescent vesicles of the size of LCM are found inside cells, suggesting that LCM remain as droplets during the internalization process, a fact consistent with endocytosis. The time course of internalization and the temperature dependency of the process further support this notion. [Also,] Endocytosis appears to be the main mechanism by which liposomes are internalized [32, 33,39]. ... In rat kidney cells, endocytotic vesicles eventually fuse to generate lysosomes [40]. If this were the case for LCM, one would expect to see the appearance, with time, of large acidic vesicles. The formation of fluorescently labeled aggregates, and the coincidence of some of them with acidic compartments, was observed. These data suggest that when grown in culture, C6 and 9L tumor cells take up LCM via the endocytic pathway” [6].
The LCM patents [5,10,41-48] employ a technology, i.e. lipid-coated microbubbles, which is patentably distinguishable over liposome technology in several major respects:
Firstly, only nonionic (and non-phospholipid) surface-active lipids are used in the production of research-grade LCM (cf. sect. IV & V), whereas charged phospholipids are used (exclusively and/or primarily) to form the vast majority of liposomes.
Secondly, the microbubble coating in the LCM invention is a single monolayer of lipid which faces a bubble core containing only a gas or nonpolar phase (as is required for maximum thermodynamic stability of a lipid monolayer system), whereas a liposome’s innermost structural layer is typically a lipid bilayer so that the core must always be a polar/aqueous medium.
Thirdly, liposome production routinely employs sonication to produce the liposome drug-delivery vehicle, whereas sonication is not necessary nor used for the production of research-grade LCM for drug delivery.
The history of intellectual property development, as well as the ownership of resulting patents, within the two separate fields of stealth liposomes versus LCM technology is also very different. As concerns liposome technology, many competing pharmaceutical companies and different university research groups are known to be exploring [e.g., 3,49] stealth liposome technology. In contrast, all patents worldwide covering the LCM technology (Filmix®) are owned by, and all multisite R&D on LCM medical applications is coordinated by, one company exclusively -- i.e., Cav-Con Inc. The major international markets (or countries) in which these LCM patents have issued, and their identification numbers, are as follows:
United States Patent No. 4,684,479 (issued August 4, 1987);
United States Patent No. 5,215,680 (issued June 1, 1993);
Canadian Patent No. 1,267,055 (issued March 27, 1990);
Japanese Patent No. 1,815,442 (issued January 18, 1994);
Australian Patent No. 657480 (issued July 24, 1995);
European P.O. Patent 0467031 (granted July 30, 1997).
Germany Patent No. DE69127032T2 (issued February 26, 1998).
United Kingdom Patent No. 0467031 (issued July 30, 1997).
France Patent No. 0467031 (issued July 30, 1997).
Italy Patent No. 0467031 (issued July 30, 1997).
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2. Haumann, B.F. 1995. Inform 6:793-802.
3. Fisher, D., et al. 1995. In: Proceedings of the Fourth Liposome Research Days
Conference, p. L26; Freiburg, Germany.
4. Allen, T.M., et al. 1991. Biochim. Biophys. Acta 1066:29.
5. D’Arrigo, J.S. 1987. U.S. Patent No. 4,684,479.
6. Barbarese, E., et al. 1995. J. Neuro-Oncology 26:25-34.
7. Simon, R.H., et al. 1995. In: Proceedings
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Germany.
8. Ho, S.Y., et al. 1996. Neurosurgery (submitted for publication).
9. D’Arrigo, J.S. 1986. Stable Gas-in-Liquid Emulsions: Production in Natural
Waters and Artificial Media, 220 pp.; Elsevier Science Publishers, Amsterdam.
10. D’Arrigo, J.S. 1993. U.S. Patent No. 5,215,680.
11. D’Arrigo, J.S. 1978. Aviation, Space & Environ. Med. 49:358-361.
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34. Foss, R.D., et al. 1991. J. Neuro-Oncology 11:185-197.
35. Stewart, P.A., et al. 1985. Acta Neuropathol. 67:96-102.
36. Huang, S.K., et al. 1994. Cancer Res. 54:2186-2191.
36a D'Arrigo, J.S. 2003. Stable Gas-in-Liquid Emulsions: Production in Natural Waters and Artificial Media, Second edition, 323 pp.; Elsevier Science Publishers, Amsterdam and NY.
37. Davis, M.A., et al. 1992. Med. Physics 19:1138.
38. Acarin, L., et al. 1994. J. Histochem. Cytochem. 42:1033-1041.
39. Straubinger, R.M., et al. 1983. Cell 32:1069-1079.
40. Griffiths, G., et al. 1988. Cell 52:329-341.
41. D’Arrigo, J.S. 1990. Canadian Patent No. 1,267,055.
42. D’Arrigo, J.S. 1994. Japanese Patent No. 1,815,442.
43. D’Arrigo, J.S. 1995. Australian Patent No. 657480.
44. D’Arrigo, J.S. 1997. European P.O. Patent No. 0467031 (granted July 30, 1997).
45. D'Arrigo, J.S. 1997. United Kingdom Patent No. 0467031.
46. D'Arrigo, J.S. 1997. France Patent No. 0467031.
47. D'Arrigo, J.S. 1997. Italy Patent No. 0467031.
48. D'Arrigo, J.S. 1998. Germany Patent No. DE69127032T2.
49. Lasic, D.D., & F.J. Martin, editors. 1995. Stealth Liposomes, 320 pp.;
CRC Press, Boca Raton FL.