GR1009114B - Pentablock polypeptidic and hybrid polymers to give in situ-forming injectable, self-healing, ph-and temperature-responsive hydrogels for localised targeted delivery of gemcitabine - Google Patents

Abstract

I ABSTRACT The invention relates: A. The synthesis of new polymeric materials, B. The polymeric materials form hydrogels with unique properties: is self-healable and thus becomes injectable, and additionally is responsive to pH and temperature. C. The polymeric materials form a mixture suitable for targeted delivery of gemcitabine to pancreatic cancer in solid tissue form. The composition consists of the polymeric materials, water and gemcitabine, forming strong hydrogels and transfers gemcitabine selectively to local pancreatic cancer tissue. Two polymeric materials are presented. One is a pentablock tripolymer of poly(L-lysine) and poly(L-histidine-co-y-benzyl L-glutamate), while the other one is a hybrid tetrapolymer containing poly(ethylene oxide), poly(L-lysine) and poly(L-histidine-co-v-benzyl L-glutamate). The polymeric materials are injectable, in-situ forming and rapidly self-healing, they are not cytotoxic up to 400 mg/kg after in vivo studies in mice, according to the National Cancer Institute of USA. The composition responds to temperature and pH through a conversion from liquid to solid and a transformation of the secondary structure of the polymeric material. This multifunctionality allows the mixture to be placed near and in contact with the cancer tissue with the simplest possible application, i.e. via injection with a very thin needle. Furthermore, due to pH- and temperature-responsiveness, the mixtures melt and convert to liquids near the cancer tissue and thus are absorbed selectively by it, while they remain strong hydrogels at the healthy tissue, blocking release of the drug. In this way, the release of the drug is directed only to the cancer tissue. Our invention is unique and operates differently from other pH- and temperature-responsive hydrogel systems where the conversion from viscous liquid to hydrogel is effected in situ due to the lower pH and the higher temperature of the cancer tissues. Our composition is converted in situ from hydrogel to liquid due to conversion of the secondary structure, independently from the pH and temperature, and then melts selectively only towards the cancer tissue due to the lower pH and the higher temperature, delivering and releasing the drug in a targeted and time-controlled manner.

Description

LINEAR PENTA-COMPONENT POLYPEPTIDES AND HYBRID POLYMERS FOR IN SITU-FORMABLE INJECTABLE HYDROGELS THAT CAN BE SELF-ORDERED AND PH- AND TEMPERATURE-RESISTANT FOR TOPICAL AND DIRECTED DELIVERY OF GEMCITABINE

DEPOSITORS: University of Athens (EKPA),

University of Thessaly (PTH)

Institute of Technology and Research (ITE)

Professor Ermolaos Iatrou, EKPA

Assistant Professor Konstantinos Dimas, P.Th.

Dr. Chrysiida Tsimplouli, P.Th.

Dr. Panagiotis Bilalis, EKPA

INVENTORS: Prof. Ermolaos Iatrou, EKPA

Prof. Dimitris Vlassopoulos, ITE

Assistant Professor Konstantinos Dimas, P.Th.

Dr. Chrysiida Tsibloumi, P.Th.

Dr. Panagiotis Bilalis, EKPA

DESCRIPTION

LINEAR PENTA-COMPONENT POLYPEPTIDE AND HYBRID POLYMER PA INJECTABLE HYDROGELS THAT FORM IN SITE AND ARE SELF-REGULATING AND TEMPERATURE-ABLE TO BE DISCHARGED PA TOPICAL AND DIRECTED DELIVERY OF GEMSITA

TECHNICAL FIELD OF THE INVENTION

[0001]      The invention presents two new polymeric materials that can be self-healing, self-healing and responsive to different pH and temperature values. The invention combines the two new polymers with two different compositions and different methodologies for the local administration of gemcitabine to pancreatic cancer tissues. Polymeric materials are biodegradable, non-toxic and have the ability to become liquid under the influence of low shear stress. When they become liquid, they can be inserted into the syringe and can be implanted close to the cancer tissue so that it comes into contact with it. Due to the structure of the polymeric material, when injected into the animal, it immediately turns into solid hydrogels containing drug. Because of its responsiveness to pH and temperature, the hydrogel melts only near the cancerous tissue, thereby selectively and directionally releasing the drug to the cancerous rather than the healthy tissue. Linear pentablock polymers can be used for the controlled release of different drugs and genes into tissues of various types of cancer that can be treated and combated through a local rather than systemic delivery methodology.

DESCRIPTION OF THE RELEVANT TECHNOLOGY OF THE INDUSTRY

[0002] Although overall cancer mortality has been steadily decreasing in recent decades due to improvement in the effectiveness of new drugs as well as early prognosis, there are some types of cancer such as pancreatic cancer where mortality is still increasing. Although the delivery of drugs using nanoparticles as drug carriers and their intravenous administration has the potential to transport drugs to most parts of the body, a very efficient technology is needed for the focused release of drugs and the bioaccumulation of these nanoparticles selectively in cancerous tissues instead of healthy tissues. In addition, there are significant biological barriers that prevent nanoparticles from reaching some target tissues such as pancreatic cancer (PC). KP has the unique property, among cancer types, of a high ratio of stromal tissue to epithelial tissue. The cancerous layer increases the pressure of the fluids inside the cancerous tissue, consequently hindering the free circulation of the nanoparticles through the blood in these tissues due to the reduced blood flow in the cancerous tissues, in relation to the healthy tissues, and thus the drugs they cannot effectively reach the target. In contrast to other solid tissues where cancer-associated fibroblasts aid tumor tissue growth and angiogenesis, the fibromatous layer in CK interferes with the formation and function of vasculature and blood circulation, thereby reducing drug delivery. through blood circulation. This fact leads to ineffective systemic chemotherapy through the use of nanoparticle carriers (Tredan O.

et al. (2007) Drug resistance and the solid tumor microenvironment, J Natl Cane Inst, 99, 1441-1454, Olive KP. et al. (2009), Inhibition of Hedgehog Signaling Enhances Delivery of Chemotherapy in a Mouse Model of Pancreatic Cancer Science, 324, 1457-1461). It is clear that especially in this type of cancer it is absolutely necessary to apply a local administration technology instead of drug transport using nanoparticle carriers intravenously.

[0003]     The advantages of the local administration of chemotherapeutic drugs in the cancerous target tissue are multiple and aim both at the better effectiveness of the drug and at the better quality of life of the patient. Implants that transport drugs and are placed directly in the cancerous tissue have the following advantages over classical carriers that are transported intravenously through the circulatory system: 1) stabilization of the encapsulated drugs and maintenance of their antitumor effect, 2) controlled and long-term release of drugs for better their diffusion and absorption into cancer cells for many cycles of cell division and proliferation, 3) entrapment and release of non-water-soluble drugs, 4) direct transport of drugs to the target, resulting in less drug loss, 5) lower dose frequency, even and only for one time, 6) less side effects. Furthermore, long-term exposure of cancer cells to chemotherapeutics for many cell divisions has been shown to lead to greater cytotoxicity and death of these cells than a high dose at a lower frequency. Only 10-15% of cancer cells are expected to be in the mitotic phase at any given time, and so this fact reduces the sensitivity of these cells to anticancer drug carriers when given for short periods of time at high doses and then repeated she. The use of an implant that releases drugs over a long period of time is beneficial when a) the possibility of regenerating cancerous tissue cannot be ruled out for all patients by a treatment involving painful intravenous drug administration, b) the disease cannot be treated with surgery or other standard means or c) the local control of the cancerous tissue makes it possible to perform a minor surgery in patients whose condition does not allow to withstand a major major surgery.

[0004] The best systems to achieve the local release of drugs are hydrogels. However, in order to achieve efficient and selective drug delivery through hydrogels, drug-carrying materials must exhibit specific characteristics in order to achieve efficient and sustained release, and must be able to be implanted in the simplest and least painful way in order to to improve the patient's quality of life. The drug carrier must exhibit multiple functions in order to implant the drug carrier into the body without surgery to prolong their release kinetics from the hydrogels, and to expand the nature of the drugs that can be delivered.

[0005] Currently, most of the materials used to transport drugs through hydrogels only work when injected into cancer tissue, in order for the lower pH and higher temperature of the cancer tissue to lead to the in situ formation of the hydrogels, and usually leads to their non-targeted and non-selective release.

In US Patent No. US 2003/0147958, Cheol-Hee Ahn et al., synthesized multicomponent copolymers composed of polypeptides and poly(ethylene oxide) for the transport of pharmaceutical compounds as well as DNA, showing similarities to the polymeric materials presented in this patent. The difference between the present patent and the one we mentioned is that in the above US patent the carriers are nanoparticles and not hydrogels.

In United States Patent US 2008/0274190 (US 2008/0274190 A1, 11/2008, Doo Sung LEE et al., 424/486) the inventors present a hydrogel that is responsive to pH and temperature as a drug delivery system. The polymeric materials in this patent are composed of oligomers of poly(ethylene oxide) and polyurethane to form a hydrogel in contrast to the polymeric materials of polypeptides presented in the present patent. Also, the responsivity of the polymeric materials act contrary to our right invention, where the hydrogel is formed due to the secondary structure and turns into a liquid at low pH and high temperature in contrast to the patent of Doo Sung LEE KOL associates which is formed under the conditions they.

[0008]        In another United States patent US 2013/0022545 (US 2013/0022545 A1, 01/2013, Doo Sung Lee et al., 424/91), a pH- and temperature-responsive hydrogel for drug delivery to the liver. The polymeric material consists of a pH-sensitive block copolymer based on either poly([3-aminoesters), or poly(amidoamines), or poly(aminoester urethanes) or poly(aminourethane ureas) or sulfamido-block copolymers . The hydrogel responds in an opposite way to our material in the present invention. The hydrogel in the above work is formed inside the cancerous liver tissue after the polymeric material is converted from a liquid to a hydrogel at the lower pH and higher temperature of the cancerous tissue.

In United States Patent US 2014/0004185 (US 2014/0004185 A1, 01/2014, Nthato ChirWa et al., 424/455) the inventors present a polymeric material that forms a hydrogel at core body temperature, while swelling slowly responding to higher pH values of cancer cells. It is used as a drug delivery system for the treatment of many types of cancer such as ovarian, pancreatic, prostate, metastatic breast, bladder and lung cancer. Polymeric materials are composed of many amino acids but not poly(histidine). The temperature responsiveness works opposite to the present invention, i.e. it forms a hydrogel at the temperature of the cancerous tissue while in our invention it swells at the temperature of the tumor releasing the chemotherapeutic drug and is based on poly(histidine).

In another invention of the United States numbered US 8,586,087 B2 (US 8,586,087 B2, 11/2013, Doo Sung Lee et al., 424/487), an amphiphilic pentacomponent copolymer that is responsive to temperature and pH as also a hydrophilic thermoresponsive tricomponent copolymer to form a responsive drug delivery system. Polymeric materials are completely different from our invention. They consist of either poly(ethylene glycol)-based materials, or poly(lactides), poly(glycolides), poly(caprolactone) random copolymer poly(caprolactone-co-lactides), random copolymer poly(caprolactone-co-glycolides), random copolymer poly( lactides-co-glycolides) and mixtures of the above, in combination with oligomers based on pH-responsive polyurethanes. These polymers include diisocyanates that have been polymerized with a diol that includes a tertiary amine in the main chain that is one or more of the amines 1,3-bis{1-(2-hydroxyethyl)-4-pyridyl}propane and 1,4- bis(2-hydroxyethyl)piperazine. In this hydrogel the responsiveness to temperature and pH works in such a way that the hydrogel forms under cancer tissue conditions, unlike our invention which swells and melts under those conditions.

In a European patent number EP 1 386 935 A1 (EP 1 386 935 A1, 06/2004, Kim Jungahn et al., C08F 293/00), the inventors presented a pH-temperature responsive system for the formation hydrogel in situ. As in all the above cases, the polymers used respond in such a way that the polymer material with water and the drug is transformed from a viscous liquid to a strong hydrogel due to the difference in pH and temperature of the cancer tissue. In our invention the mixture of polymer, water and drug forms a strong hydrogel due to the structure of the polymer and its secondary structure, and then it swells and degrades selectively near the cancerous tissue achieving the release of the drug selectively in the cancerous and not in the healthy.

[0012] Similar pH-responsive hydrogels have been presented by Yi-Fang Zeng et al. (Yi-Fang Zeng et al. (2012) Controlled delivery of recombinant adeno-associated virus serotype 2 using pH-sensitive poly(ethylene glycol)-poly-L-histidine hydrogels, Biomaterials 33 9239-9245). The polymers used were bicomponent copolymers of poly(ethylene oxide) and poly(histidine). The hydrogels were used to transfer the green fluorescent protein gene rAAV2-GFP. By lowering the pH, the poly(histidine) cluster that forms the hydrogel swells thereby releasing the protein.

Finally, hydrogels composed exclusively of polypeptides were synthesized by Zhibo Li et al. (Zhibo Li et al. (2010). Tunable hydrogel morphology via self-assembly of amphiphilic pentablock copolypeptides, Soft Matter, 6, 2546-2551) using pentacomponents amphiphilic copolypeptides. The pentameric polypeptides consisted of poly(L-lysine) and poly(L-leucine). The polymers were neither pH nor temperature responsive and were not used for drug delivery applications.

[0014] In all cases mentioned the inventors presented a pH- and temperature-responsive system that forms a hydrogel inside the cancer tissue due to the higher temperature and lower pH of this tissue (Antoni Sanchez-Ferrer et al. (2013), Thermo -responsive peptide-based triblock copolymer hydrogels, Soft Matter, 9, 4304, Narendra Singh et al. (2014), In situ gelling pH- and temperature-sensitive biodegradable block copolymer hydrogels for drug delivery, Journal of Controlled Release, 193, 214 -227, Chaoliang He et al. (2008), In situ gelling stimuli-sensitive block copolymer hydrogels for drug delivery. Journal of Controlled Release 127, 189-207). But injecting such a large amount of solid material into the cancerous tissue will result in tissue injury and the creation of metastatic cancerous tissues. Although there are a large number of commercially available hydrogels that carry drugs to fight eye diseases or as wound dressings to fight chronic diseases, to our knowledge there are no commercially available "smart" hydrogels for the local delivery of chemotherapeutic drugs to fight cancer that to exhibit the necessary properties mentioned. Advanced synthetic chemical methodologies are required to produce polymeric materials that will exhibit the multifunctionality of a "smart" hydrogel. Golinska M. et al. presented a tricomponent tetrapolymer that was pH-responsive and formed self-healing hydrogels, but they were not used for drug delivery to cancer tissues (Golinska M. et al. (2014) Dilute Self-Healing Hydrogels of Silk-Collagen- Like Block Copolypeptides at Neutral pH, Biomacromolecules, 15, 699-706).

OUTLINE OF THE FIGURES

[0015] FIG. 1. Schematic representation of the invention and the mechanism of the target

transfer of Gemcitabine (gemcitabine).

[0016] FIG. 2. The polymeric structures of the invention used in the formulation

hydrogels.

[0017] FIG. 3. Reactions used to synthesize pentacomponent terpolymers

that form the hydrogels (PTY).

[0018] FIG. 4. Size exclusion chromatography for the synthesis of PTY polymeric material with a mobile phase of water/acetonitrile/trifluoroacetic acid in a composition by volume of 60/39.9/0.1 (v.v). From right to left the peaks: poly(L-lysine) difunctional, poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine)-δ-poly(L-histidine -co-γ-benzyl ester of L-glutamic acid), poly(L-lysine)-d-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine)-b -poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine). It is clear that the polymers have a small molecular weight distribution and a high homogeneity with respect to

molecular weight and composition.

[0019] FIG. 5. Cyclic dichroism of PTY at two pH values.

[0020] FIG. 6. Rheology Measurements at Various Concentrations. It is obvious that its power

of hydrogel depends on the concn.

[0021] FIG. 7. FIG. 7A shows the pH dependence of the PTY hydrogel at various pH values, namely the pH values of healthy tissue, cancer tissue, and final endosomal pH. It is clear that the hydrogel turns into a liquid going from the healthy tissue to the cancerous one. FIG.

7B shows that the PTY hydrogel has almost zero temperature responsiveness.

[0022] FIG. 8. Schematic representation of the mechanism that makes the hydrogel self-repair almost instantaneous: the pentablock polymer structure and the presence of two hydrophobic groups in a polymer chain creates a memory in the structure and formation of the hydrogel after

injected instantly.

[0023] FIG. 9. The time required for the void hydrogel to form after the application of a strong shear stress that turns the hydrogel into a liquid: only a few seconds are required to return to the initial G' value when the shear stress is stopped.

[0024] FIG. 10. pH dependence of gemcitabine-encapsulated PTY between values of healthy and cancerous tissue. The strength of the hydrogel decreases significantly at the pH of the tumor

tissue.

[0025] FIG. 11. Recovery of hydrogel strength G' and G" (self-correction) as a function of blank hydrogel concentration at 37 °C and pH=7.4.

[0026] FIG. 12. Restoring hydrogel strength G' and G" (self-healing) of blank hydrogel at 37 °C.

[0027] FIG. 13. Recovery of hydrogel strength G' and G" (self-healing) as a function of pH for the hydrogel with entrapped drug gemcitabine at a drug:polymer ratio equal to 1:1 (w/w) at pH=7.4 and T=37 ° C.

[0028] FIG. 14. Vials containing the same amount of polymer and water at different pH values. Left: pH=6.0, center: pH=7.4, right: pH=6.5. It is evident that at pH=7.4 it is the strongest hydrogel, while at pH=6.0 the mixture becomes almost liquid, and at pH=6.5 the material is in an intermediate state.

[0029] FIG. 15. SEM (Scanning Electron Microscopy) micrographs of the hydrogel at pH=6.5 (left image), and at pH=6.5 (right image).

[0030] FIG. 16. Reactions used to synthesize the pH- and temperature-responsive hydrogel-forming pentacomponent tetrapolymer (APTY).

[0031] FIG. 17. Rheology measurements of the empty APTY hydrogel at 25 °C as a function of pH, at pH=7.4 (dark triangle) and 6.5 (light triangle). It is clear that the hydrogel responds to pH values

between those of healthy and cancerous tissue.

[0032] FIG. 18. Rheological measurements of the APTY gap at 37 °C as a function of pH, at pH=7.4 (dark triangle) and 6.5 (light triangle). It is evident that hydrogels respond to pH values

between those of healthy and cancerous tissue

[0033] FIG. 19. Rheological measurements of the APTY gap at 41 °C (cancer tissue temperature) as a function of pH, at pH=7.4 (dark triangle) and 6.5 (light triangle). It is evident that the hydrogel responds to pH values between healthy and cancerous tissue.

[0034] FIG. 20. Rheological measurements of APTY blank at pH=7.4 as a function of temperature, at 25 °C (intermediate dim), 37 °C (dimest) and 40 (darkest). It is evident that the hydrogel responds to temperature at values between those of healthy and cancerous tissue.

[0035] FIG. 21. In vivo toxicity test of the empty pentameric poly(L-lysine)-bpoly(L-histidine-co-L-leucine)-b-poly(L-lysine)-b-poly(L-histidine-co-L -leucine)-b-poly(L-lysine). From left to right: 400 mg/Kg, 200 mg/kg, 100 mg/Kg.

[0036] FIG. 22. In vivo toxicity test of the empty pentameric poly(L-lysine)-bpoly(L-histidine-co-y-benzyl-L-glutamate)-b-poly(L-lysine)-b-poly(L-histidine -co-γ-benzyl-L-glutamate)-bpoly(L-lysine). From left to right: 400 mg/Kg, 200 mg/kg, 100 mg/Kg.

[0037] FIG. 23. Tumor tissue size as a function of time for different injection groups in mice with human pancreatic cancer xenografts. A. Without any administration, B. Hydrogel, i.e. injection of the same amount of empty hydrogel as the hydrogel animals received with the entrapped drug, C. Intraperitoneal injection of pure gemcitabine, D. Subcutaneous injection of pure gemcitabine close to the tumor tissue, E .Injection of the hydrogel with the encapsulated drug Hydrogem, close to the cancerous tissue.

[0038] FIG. 24. Gemcitabine in hydrogel with only one dose (dpi 17) shows %DT/DC superior to systemic gemcitabine (two doses dpi 17&24) and slightly better than topical administration of

of free drug (two doses-dpi 17&24).

[0039] FIG. 25. Kaplan-Meier survival curves of different injection regimens in models

human pancreatic cancer mice.

[0040] FIG. 26. The hydrogel prefers to adhere to cancerous tissue over healthy tissue.

DESCRIPTION

[0041] The present inventors present formulations of either synthetic amphiphilic penta-component hydrogel-forming terpolymers (PTY), wherein the polymeric material is of the type poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L -glutamic acid)-b-poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly( L-lysine ) (Plys-b-(PHIS- co-PBLG)— b-PLys-b-(PHlS-co-PBLG)-b-PlYs) or pentacomponent hydrogel-forming tetrapolymers (APTY), where the polymeric material is of the type poly (L-lysine )-b-poly (L-histidine-co-γ-benzyl-L-glutamate)-δ-poly(ethylene oxide)-δ-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine). Polymers are suitable for selective local delivery of gemcitabine to pancreatic cancer and not to healthy tissues. The inventors present methods of synthesizing these PTY and APTY and show that both can be very easily implanted by injection through a fine needle (type 21G) and can be directly re-formed near cancerous tissues due to the combination of their advanced macromolecular architecture and the unique properties of the hydrophobic part of polymers whose secondary structure is destroyed under shear stress and quickly self-repairs when the shear stress stops (Ashton et al. (2010), Susceptibility of Different Proteins to Flow-Induced Conformational Changes Monitored with Raman Spectroscopy Biophysical Journal, 98, 707-714, and Bekard et al. (2010), α-Helix unfolding in simple shear flow. Soft Matter, 2011, 7, 203-209). Once the hydrogels are injected near cancerous tissues, they can selectively release gemcitabine into cancer rather than healthy tissues due to their responsiveness to pH and temperature. We take advantage of the lower pH and higher temperature of cancerous tissue compared to healthy tissues to induce melting of the hydrogel only near the cancerous tissue, directing drug delivery to the cancer and not to the healthy tissues.

Almost all pH and temperature responsive hydrogels presented so far in the literature or patents are based on their ability to transform from a viscous liquid at room temperature and healthy tissue pH (value 7.2 to 7.6 by average) in the form of a hydrogel that when injected into the cancer tissue, takes advantage of the higher temperature and/or lower pH of the cancer tissue. Thus, the difference in pH and temperature is used to convert the injectable material from liquid to hydrogel. Our innovative hydrogels work completely differently from the inventions presented so far. PTY and APTY have two different functions. One is the ability to make liquids under low shear stress to be introduced into the syringe by simple suction, and immediately become hydrogels when injected out of the needle. This allows easy and controlled implantation of the hydrogel close to the cancerous tissue, rather than inside it. Therefore, just by stirring the needle inside the vial of hydrogel it becomes a liquid that allows its uptake into the syringe (where in the form of hydrogel this is not possible) and when injected close to the cancerous tissue it is reconstituted into a hydrogel in less than 7 seconds, not because of their pH and temperature response as in the hydrogels presented so far, but because of its macromolecular architecture and the secondary structure of the hydrophobic component. Then, due to the multi-functionality of the hydrogels, the formed hydrogel melts only near and in contact with the cancer tissue, due to its pH and temperature responsiveness. This operation is in stark contrast to the systems presented so far, where the lower pH and higher temperature of the hydrogel cause the formation of a hydrogel form instead of the formation of a liquid, as in our invention. In addition to gemcitabine, other therapeutic molecules that can be selectively delivered by hydrogels are general chemotherapeutic agents and DNA molecules. PTY and APTY can be used for topical administration in other cancers, such as ovarian and colon.

Addition of the drug to the polymer composition with water does not affect the shear stress properties of the polymer. Therefore, the polymer composition with water and drug quickly self-heals after applying the shear rate, as observed for the polymer-water composition without drug.

[0044] The advantages of our invention regarding its functionality are the following (FIG. 1):

1.   It requires the simplest, least invasive possible implantation, ie not by surgery, but by an injectable hydrogel-forming epitope that will be placed in the area of the cancerous tissue, not in it.

2. It exhibits slow release characteristics. Injection of the hydrogel with the drug(s) and/or genes will be done every month or even later.

3. Targeted release of drug and/or genes to cancer tissues reduces cytotoxic side effects in healthy organs.

4. It has the ability to transport a mixture of already approved chemotherapeutic drugs (hydrophilic gemcitabine and hydrophobic paclitaxel) regardless of their solubility in water.

5. The material is non-cytotoxic and biodegradable, so surgical removal is not necessary after drug implantation.

[0045] Other advantages of PTY and APTY include their safety, their efficacy, the ability to manufacture them according to GMP, and the ability to increase their production to industrial production levels. PTY as well as APTY can be easily produced in high purity and in large quantities, and the chemicals used in their synthesis can be completely removed. In addition, the samples are free of pyrogens or biological contamination. By adjusting the pH and polymer concentration of the hydrogels their potency can be increased to desired levels, and thus the timing of targeted drug delivery can be controlled. The unprecedented flexibility and properties of these pentacomponent hydrogels provide multiple molecular tailoring to tune their different properties and achieve the desired functionality, such as sustained sustained release, amount of drug entrapped, transport of either hydrophilic or hydrophobic drugs, combined transport of hydrophilic and hydrophobic drugs, as well as the transport of drug-entrapped nanoparticles.

[0046]    By combining careful precision synthesis and multifunctional materials we have synthesized a material that self-organizes hierarchically from a few Angstroms to micrometers to achieve the desired properties. The polymeric material forms an injectable, in situ hydrogel through physical interactions. The hydrogel is converted into a fluid liquid by applying slight shear stress through agitation with a syringe needle into its volume in order to be drawn up by the syringe, and when injected near the cancerous tissue it becomes a hydrogel within seconds, regardless of the characteristics of the tissue. The material is sensitive to pH and leads to its melting at pH values characteristic of cancerous tissue, i.e. pH = 6.5 (Leo E. Gerweck et al. (1996), Cellular pH Gradient in Tumor versus Normal Tissue: Potential Exploitation for the Treatment of Cancer, Cancer Research 56, 1194-1198), while remaining in the form of a hydrogel under the conditions of healthy tissue. The polymeric composition of the hydrogel contains a larger amount of water than that of the cancerous tissue. Due to the osmotic pressure, the material is absorbed together with the drug into the cancer tissue thus achieving the selective and efficient transport of the entrapped chemotherapeutic agent. The release of the drug is limited and directed towards the healthy tissue since it remains in the form of a hydrogel, and it is known that the diffusion of water and drugs through a hydrogel can be significantly slower than in a nanoparticle, depending on the system. The hydrogel has the ability to selectively adhere to cancerous tissue through electrostatic interactions in order to maintain a constant rate of degradation. In addition, it has a buffering capacity, in order to remain in the form of a hydrogel for a significant period of time and to be released slowly.

[0047] The functionality of the hydrogel is not only limited to the targeted delivery of the drug selectively to cancer tissues, but also helps in efficient and quantitative delivery into the cytoplasm of cancer cells. The polymer incorporated together with the drug into the cell by endocytosis has the ability to breach the endosome, and consequently the drug is released into the cytoplasm, avoiding exocytosis and chemoresistance, (Polymeric Gene Delivery, Edited by M. Amiji, Chapter 13, p246-261, Nelson N. Structure and Pharmacology of proton-ATPases, Trends Pharmacol Sci 12, 71, 1991).

The macromolecular architecture, which gives the desired possibility of in situ formation of an injectable hydrogel with rapid self-healing properties, is a five-component polymer consisting of a middle hydrophilic block, two hydrophobic blocks, followed by two hydrophilic segments, each attached to the other side of the hydrophobic clusters. In order to achieve the desired properties, we synthesized two different penta-component polymers, a tripolypeptide and a hybrid tetrapolypeptide. The chemical structures of the polymers are given in FIG. 2.

Initially, the hydrophobic parts of the pentacomponents were very (L-histidine-co-L-leucine) with a molecular ratio of 50% for each amino acid, L-histidine and L-leucine. However, in vivo acute toxicity tests of the polymer formulation with drug-free water showed that the hydrogel did not show cytotoxicity up to 100 mg/Kg of mice, while for concentrations of 200 and 400 mg/kg it was toxic. Consequently, poly(γ-benzyl ester of -L-glutamic acid) (PBLG) is used instead of poly(L-leucine) (PLEU), in the same molar ratio of 50% since the blank hydrogel experiment did not show any cytotoxicity in the acute tests toxicity for concentrations up to 400 mg/Kg, as accepted by the US National Cancer Institute (NCI).

The final materials used for the invention are two. One material is the pentameric tripolypeptide of the formula ABA'BA where A and A'' are poly(L-lysine) of different molecular weights, and B is poly(L-histidine-co-γ-benzyl ester of L-glutamate acid). The second material is similar to that described above, with the only difference that the middle block A' is very (ethylene oxide) (PEO) and not very (L-lysine).

EXAMPLES

Example 1.

Synthesis of amphiphilic pentameric tripolypeptides that form hydrogels

(PTY)

A. Summary.

[0051] Amphiphilic pentameric tripolypeptides that form hydrogels (PTY) are synthetic materials formed by using a polymer material of the type poly(L-lysine)-bpoly(L-histidine-co-γ-benzyl ester of L-glutamic acid)- b-poly(L-lysine)-δ-poly(L-histidine-γ-benzyl ester of L-glutamic acid)-b-poly ( L-lysine ) (poly(γ-benzyl ester of L-glutamic acid): PBLG and poly(L-histidine):PHIS). Their properties can be altered by changing the PBLG/PHIS molar ratio of the hydrophobic component and can be used for biomedical applications. The size, hardness and density of PTY can be adjusted by varying the concentration of PTY (in water), pH and composition of polymeric materials. For each PTY, increased concentration resulted in increased hydrogel strength in vitro and greater viability in vivo. In addition, higher pH values also resulted in increased hydrogel strength in vitro as well as additional viability in vivo. It was found that the middle hydrophilic poly(l-lysine) block should have a high molecular weight in order to form a strong hydrogel at concentrations as low as 3% (w/w) in water. To determine the effect of PBLG composition on hydrogel properties, we synthesized a series of pentablock terpolymers where OL poly(L-lysine) blocks were always the same, while the molar ratios of PBLG/PHis PBLG varied between 70, 50, 30 and 0. It was found that the molar ratio of 50% gave the best results regarding strong pH response combined with low polymer concentration to achieve the desired hydrogel strength.

B. Methods

B.1. Composition of PTY

[0052]     a. Purification of reagents and synthesis of monomers. The purification of the solvents, as well as the synthesis of the monomers, are given in the literature (Mavrogiorgis et al. (2014), Controlled polymerization of histidine and synthesis of well-defined stimuli responsive polymers. Elucidation of the structure-aggregation relationship of this highly multifunctional material, Polym. Chem., 5, 6256-6278) Size exclusion chromatography. Size exclusion chromatography (SEC) was used to determine Mn and Mw / Mn values. Analysis was performed on a Waters Breeze instrument equipped with a 2410 differential refractometer and a Precision PD 2020 (150, 900) dual angle light scattering detector. The carrier solvent was a solution of 0.10% trifluoroacetate water/acetonitrile (60/40 v / v) at a flow rate of 0.8 mL/min at 35 °C. Three linear Waters hydrogel columns were used. Spectroscopy<1>NMR (300 MHz) was performed using a Varian Unity Plus 300/54 spectrometer. Polymer spectra were performed in either D 2 O with DCI, deuterated DMSO or CF 3 COOD, while monomer spectra were taken in CDCl 3 at room temperature. FTIR measurements were performed with a Perkin Elmer Spectrum One instrument, on KBr pellets at room temperature, in the range 450-4000 cm<-1>. Circular Dichroism was done with a Jasco J-815 model, which has a Peltier PTC-423S/ 15 model with a thermostat system. The cell used was 1 mm Quartz Suprasil. Typical concentrations were around 3x10<-4>g/ml. UV spectroscopy was performed with a Perkin Elmer Lamda 650 spectrometer from 190-500 nm, at room temperature with a cuvette requiring 120 µl.

[0053]    B. General polypeptide synthesis. The reactions used to synthesize the PTY materials are given in FIG. 3A and FIG. 3B. The polymerization reactors were designed to have a volume at least three times the volume of CO2 produced by each polymerization. Polymerizations were carried out with 1,6-hexamethylenediamine as initiator. To a solution of 1.2 g of ε-tert-butyloxycarbonyl-L-lysine N-carboxy anhydride (Lyx-NCA) in 20 ml of DMF, 2 ml of a solution of hexamethylenediamine in DMF (9 x 10<-6>mol / ml) was added . After completion of polymerization (1 day) (middle cluster), 0.54 g Nim-trityl protected L-histidine N-carboxy anhydride (HIS-NCA) and 0.34 g γ-benzyl-L-glutamate N-carboxy anhydride (BLG-NCA) (molar ratio 1:1) were dissolved in 10 ml DMF and added to the polymerization flask. Polymerization is allowed to continue for 6 days (two hydrophobic clusters). Then 0.72 g Lys-NCA in 10 ml DMF was added. After complete consumption of Lys-NCA (2 days), the pentapolymer was precipitated in 250 ml of diethyl ether, filtered and dried to constant weight (2 g). To deprotect the tertyl and tert-butyloxycarbonyl groups from the poly (histidine) and poly (L-lysine) moieties respectively, the pentapolymer was dissolved in 20 ml of dichloromethane (DCM) and an equal volume of trifluoroacetic acid (TFA) was added. The polymer was completely dissolved and allowed to deprotect for 1 hour at room temperature, and the solution turned yellow-light brown. Next, 0.2 ml of triethyl-silane was added to bind the protecting groups, and the solution from yellow-light brown became colorless. The solution was poured into 300 ml of diethyl ether, and the white solid was filtered and dried. Then, the deprotected pentapolymer was dissolved in 100 ml of DI water and then dialyzed against HCl solution (pH = 3) for 2 days, NaOH solution (pH = 9) for 2 days, and DI water for 2 days. Finally, the solution was lyophilized to obtain the final pentapolymer as a white powder (1.5 g).

B 2. Rheology.

[0054] a. Rheology of hydrogels. Pentameric terpolymers were prepared by dissolving the lyophilized polymer samples in MilliQ water.

B. Rheology of gemcitabine hydrogels. Gemcitabine-loaded pentacomponent hydrogel-forming terpolymers were prepared by first dissolving the appropriate amount of gemcitabine in MilliO water, adjusting pH = 7.4 or another pH with the required amount of 0.1 N NaOH solution in MilliO water, and the gemcitabine solution added to the polymer. Then (gemcitabine hydrogel is called hydrogem) it was left for 2-3 hours under dissolution. Rheological properties were not affected by stirring, identical mechanical properties were obtained by allowing the terpolymer to dissolve to form the hydrogel after three days without stirring.

Rheology measurements were performed using the controlled stress rheometer (ARES from Rheometric Scientific) with a transducer rebalancing force of 0.004 at 200 torque. Cone and plate geometry, stainless steel, diameter 25 mm, cone angle 0.04 rad. Sample volume 0.3 ml. Seal sample with 5 cp silicone oil around the sample to eliminate evaporation. Temperature control through recirculating water/ethylene glycol bath.

C. RESULTS

C. 1. Physicochemical Characterization of PTY.

The pentameric terpolymer was extensively characterized by a combination of methods. The molecular weights and molecular weight distribution of the final pentameric terpolymer were found by Size Exclusion Chromatography. The completion of the polymerization was monitored by Infrared Spectroscopy, through the peaks of the NCA monomers at 1820 and 1780 wavelengths. The composition was confirmed by <1>H NMR Spectroscopy. The combined Characterization methods demonstrate that the pentameric terpolymer was what was expected from stoichiometry and design during synthesis. The Exclusion Chromatograms of the PTY composition are given in FIG. 4. The molecular characteristics of PTY terpolymers are given in Table 1.

Table 1. Molecular Characteristics of Pentacomponent Tetrapolymers PLys60-b - (PHis20-co-

PBLG20)-PLys'200-b-(PHis20-co-PBLG20)-b-PLys60

C.2. Secondary Structure of Pentacomponent Tripolymers.

[0058] The Secondary structure of the terpolymers at various pH and temperatures was determined by the Cyclic Dichroism method. (FIG. 5). We studied the change of the secondary structure of the hydrophobic part of the pentapolymer by subtracting the random coil spectrum of poly(l-lysine) according to the composition of the pentapolymer. Thus, the final secondary structure is due only to the hydrophobic part that causes the organization and formation of the hydrogel. The percentage of secondary structure was calculated using Dichroplot. We found that at pH=7.4, the hydrophobic part of the polypeptide has a significant amount (60%) of β-sheet, while also having a 10% α-helix and 20% random coil. At pH=6.6, the β-sheet becomes a random coil as well as an α-helix. The β-sheet is reduced to 25%, while the helix is increased to 45% and the random glomerulus to 30%.

[0059] It is known that in order to form a strong hydrogel with a high modulus of elasticity, it is necessary to form hydrogen bonds through β-sheets (Akihiro Abe et al. (2013) Controlled Polymerization and Polymeric Structures, Springer International Publishing Switzerland, p. 178-181.). At pH=7.4 the strong hydrogel is formed due to the large amount of β-sheet. At lower pH=6.6, where polyhistidine is partially protonated, the conformation of polyhistidine changes to a mixture of α-helix and random coil. When the pH becomes even lower, the structure becomes a random coil. This change in structure reduces the modulus of elasticity of the hydrogel, and therefore the hydrogel becomes liquid, directing the transport of the drug to the cancer tissue, which has a lower pH.

C.3. Rheology measurements to study hydrogel strength as a function of pH, concentration and temperature.

Hydrogel strength as a function of water content, drug content, pH and temperature between cancer and healthy tissue values was studied by Rheology measurements. The most important hydrogel property found by Rheology was shear thinning, where by inducing a shear stress, the hydrogel became a fluid fluid, and when the shear stress stopped the hydrogel re-formed very quickly. This was a very important property of the hydrogel to be injectable, self-healing and formed in situ in the mouse body. To get the simplest possible dose of the hydrogel into the animal, we needed to modify the molecular characteristics of the polymer so that it would easily liquefy with little shear stress while being able to reassemble itself and form a hydrogel again after a few seconds. close to the cancerous tissue and not diffuse peripherally when injected as a liquid. For this purpose we had to find the dependence of the self-healing (self-repair) time as well as the strength of the hydrogel required to remain in this state under the pressure of the mouse tissues.

C.3.A. Rheological behavior of empty Hydrogels.

[0061] We first examined the rheological properties of empty hydrogels, ie the hydrogel without the drug. The parameters that affect the strength of the hydrogel are the concentration of the polymer, i.e. the percentage of polymer to water, temperature and pH. First, we studied the dependence of the potency of the empty hydrogels as a function of polymer concentration and pH at a constant temperature of 37 °C i.e. the temperature of the animal's body. It was found that the polymer concentration significantly affects the modulus of elasticity as well as the viscosity modulus of the hydrogel as shown in the Dynamic Frequency Scan as well as in the Dynamic Voltage Scan measurements of blank PTY hydrogels at pH=7.4 as shown in Figures 6A and 6B respectively.

Next, we studied the influence of pH and temperature on the empty hydrogels (Figures 7A and B). The polymer concentration in both images is 3.33% (weight by weight). Dynamic Frequency Scan plots of PTY at 37 °C are shown in Figure 7A, each at three different pH values, namely 7.4, 6.5 and 6.0. The results showed that the influence of pH is particularly very strong. The polymer turns from a strong hydrogel at the pH of healthy tissue to a very weak hydrogel, almost liquid, at the pH of the cancerous tissue, due to the parallel decrease in the elastic modulus together with the decrease in the G'/G" ratio at pH=6.5. change is very pronounced and the hydrogel becomes liquid at an even lower pH=6.0.Although the pH dependence was very strong, temperature did not affect the strength of the polymer at each pH (Figure 7B) in the temperature range between 37 °C ( healthy cell) and 41°C (cancerous tissue).

[0063] Finally, we studied the time required for the self-healing of the hydrogel, that is, the time required for the hydrogel to have its initial strength after applying a high shear stress to become liquid. Self-healing is based on the change of 3D secondary structure from α-helix or β-sheet to random coil. Shear stress, which depends on duration and frequency, can cause the secondary structure to unfold, resulting in a reduction in protein organization. As for hydrogel organization the modulus value increases with increasing β-sheet percentage, the shear stress causes the organization to be destroyed, turning the material into a flowing liquid. In any case, when the shear stress stops, the original secondary organization is restored, and the hydrogel forms surprisingly quickly. It was very important to find the appropriate macromolecular architecture to have very fast self-healing (self-repair) so that when injected, the fluid mixture of hydrogel and drug quickly forms a hydrogel and remains in the area of the cancerous tissue, rather than remaining in the liquid phase and spread into the tissue. We found that the linear pentad has the best architecture for this property. The presence of external poly(L-lysine) forces the hydrophobic chains to organize better, causing the formation of a strong hydrogel. In addition, it was necessary to use two hydrophobic parts in the same molecule, which causes the organization and formation of a strong hydrogel. The two hydrophobic components create bridges between different hydrophobic regions, resulting in the formation of a well-defined three-dimensional interconnected nanostructure surrounded by large amounts of water that maintain the polymer-water mixture in the hydrogel form. In addition, it was found that the pentameric molecular architecture was very important for the rapid self-healing (self-repair) of the hydrogel and recovery from the liquid phase. The simple dibasic terpolymer of the type poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of-L-glutamic acid) or of the tribasic poly(L-lysine)-b-poly(L-histidine -co-γ-benzyl ester of (L-glutamic acid))-b-poly(L-lysine) had the property of self-healing at significantly longer times than the pentamer terpolymer. When shear stress is applied, it will apparently cause unfolding of one of the two hydrophobic chains, destroying the cross-links and thus destroying the organization (FIG. 8). The second hydrophobic part that remains connected to other hydrophobic chains creates a "memory" of the organization, resulting in very rapid formation of the self-healing hydrogel. It was found that the pentablock terpolymer required less than 7 seconds to self-heal while the triblock and diblock required more than 50 seconds (FIG. 9).

C.3.B. Rheological behavior of hydrogel with entrapped drug (Hydrogem).

Gemcitabine is a hydrophilic drug and remains in the aqueous phase, and therefore its addition is not expected to significantly alter the rheological behavior of the hydrogel, because its stability depends on the hydrophobic part of the polymer. The most important parameter we had to perfect was the rate of drug release. Considering that the diffusion of drug and water is very slow in a hydrogel, drug release depends on the rate at which the hydrogel changes from a gel form to a liquid to achieve targeted release to the cancerous tissue. We found that the rate of this transformation depends on the initial elastic modulus of the hydrogel, which can be controlled by the polymer concentration. In order to directly correlate the rheology properties of the hydrogel with the in vivo properties of the stability and duration of the hydrogel in the body of mice, we performed a "calibration" with a subcutaneous injection in healthy mice of the same polymer with different concentrations and therefore different modulus of elasticity G', as well as modulus of viscosity G". It was found that at concentrations of 1:30 when the modulus of elasticity at 1 rad / sec was 975.21 Pa the hydrogel stayed for 36 days, when it was 830.25 Pa it stayed for 8 days, when it was 346 Pa it stayed for 2 days, while at a lower modulus of the hydrogel it did not stay in gel form, but under the body pressure of mice it stays in liquid form. The data are shown in Table 2. It is obvious that the modulus of elasticity dramatically affects the time the hydrogel stays in the mice.

Table 2. Influence of G' on the time the hydrogel turns into a liquid.

A slightly different behavior was observed when drug was entrapped in the hydrogel. We performed rheology measurements of the hydrogel with entrapped gemcitabine at a polymer concentration namely 1:30 (polymer:water (w/w)) and a drug concentration of 1:30 (drug:water (w/w)). Consequently, the drug:polymer ratio was 1:1.

[0066] It was found that the presence of the drug led to a slight stabilization of the hydrogel at pH = 7.4 at 37 °C indicating the higher elastic modulus compared to that of the blank (FIG. 10A). The modulus of elasticity increased from 975 for the hydrogel to 1100 Pa for hydrogem (gemcitabine hydrogel) at this pH and temperature. At pH = 6.5 and temperature 37 °C, the modulus reduction is not as high as in the case of the empty hydrogel at the same temperature and pH. However, since the decrease in the modulus of the empty polymer at pH = 6.5 was too large, this will lead to a rapid collapse of the hydrogel and lose its slow-release properties. Consequently, the reduction to 350 Pa that occurs in the hydrogem is marginal and will maintain the slow conversion of the hydrogel to liquid and hence the slow release of the drug.

[0067] At 40 °C we see a slightly different behavior at low frequencies (FIG. 10B). In the case of the blank hydrogel we did not see any significant difference in either the modulus of elasticity or the modulus of viscosity compared to the values at 37 °C. At 40 °C, the modulus of elasticity is almost identical to that obtained at 37 °C, however the viscosity modulus is greater at low frequencies, which allows the hydrojem to turn into liquid faster at 40 °C (tumor tissue) and not at 37 °C (healthy tissue). Therefore, it is clear that in addition to the pH response there is also a weak temperature response of the hydrogem.

C.3.C. Study of Self-Healing (self-healing) properties as a function of polymer concentration and pH.

[0068] The self-healing (shear thinning) properties of the blank hydrogel and hydrogem (gemcitabine-entrapped hydrogel) were carefully exploited because these properties are necessary to make the hydrogel injectable. First, the self-healing of the hollow hydrogel was studied by rheology. A strong shear stress was applied to the hydrogels and hydrogems, making them liquid. The shear stress (time 0) was then stopped and the time required for them to return to exactly the same G value was measured. FIG. 13) to self-heal.

[0069] It is evident that in all cases (FIGS. 11-13) the time required for the hydrogel to self-heal is less than 7 seconds. This means that during the injection of each hydrogel (hydrogem) formulation into the mice, even before the completion of the injection the hydrogel has recovered and self-healed. This is due to the unfolding of the secondary structure induced during the application of the shear stress leading to a partial destruction of the self-organization, thus creating a viscous liquid. However, due to the existence of two aggregation moieties per single chain, even if one of the two segments unwinds the other probably continues to aggregate creating a memory for the polymer chains to quickly recover when the shear stress ceases. The macromolecular architecture creates "memory" in the molecules to quickly recover after the liquid hydrogel injection is completed in mice (shear stress state).

C.4. Scanning Electron Microscopy (SEM) of the hydrogel.

Three different vials of the same amount of polymer and water were pH-adjusted by adding a small amount of dilute HCl, i.e. prepared at pH=7.4, pH=6.5 and pH=6.0. The final mixtures are depicted in FIG. 14. Sample of the hydrogel with pH=7.4 and pH=6.5 were flash frozen with liquid nitrogen and then the water was lyophilized under high vacuum. The EMF of the samples is shown in FIG. 15. It is evident that the EMC at pH=7.4 has a three-dimensional structure of interconnected nanostructures that keeps the mixture in the form of a hydrogel, while the structure at 6.5 collapses giving the form of an almost liquid structure.

EXAMPLE 2

Synthesis of amphiphilic hybrid pentacomponent tetrapolymer hydrogels (APTY)

A. Summary.

[0071] Amphiphilic pentablock hybrid hydrogel-forming tetrapolymers (APTYs) are synthetic materials formed by using a polymer material of the type poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamate acid)-poly(ethylene oxide)-b-poly(L-histidine-co-γ-L-glutamic acid benzyl ester)-b-poly(L-lysine). Their properties can be altered by changing the poly-γ-benzyl ester of L-glutamic acid/poly(L-histidine) molar ratio of the hydrophobic part and can be used for biomedical applications. The size, stiffness and density of APTY can be tuned by varying the APTY concentration, pH and composition. For each APTY, increased concentration or higher pH values or lower temperature resulted in increased gel strength in vitro. It was found that the hydrophilic poly(ethylene oxide) middle block should have a high molecular weight in order to form strong hydrogels at concentrations as low as 4% (w/w) in water. The molecular weight of the poly(ethylene oxide) middle block ranged from 6.0 x 10<3>, 10.0 x 10<3> and 20 x 10<3>g/mol, while the (L-histidine-co-γ-benzyl ester of L -glutamic acid) and poly (L-lysine) remained stable. The molecular weight of the poly(L-lysine) cluster had a molecular weight of 7.7 x 10<3>g / mol, while the poly(L-histidine-co-γ-benzyl ester of L-glutamic acid) cluster had a molecular weight of 7.1 x 10<3>g/mol. It was found that the molar ratio of 50% by mol of poly(L-histidine) in the polypeptide cluster gave the best results regarding strong pH and temperature response in combination with low polymer concentration to achieve the desired strength of the hydrogel.

B. Methods.

B.l. Composition of APTY

The materials used for the synthesis of the polymers, as well as the characterization methods were the same as those described in Experiment 1. α,ω-Diamino terminated PEG was purchased commercially. Characterization Methods are described in Experiment 1.

[0073] General polypeptide synthesis. The reactions used to synthesize the APTY materials are shown in FIG. 16A and FIG. 16B. The polymerization reactors were designed to have a volume at least three times the volume of CO2 produced by each polymerization. Polymerizations were carried out with a,o-diamino acroreactive polyethylene oxides as a bireactive macroinitiator. For the synthesis of the pentacomponent poly(L-lysine)-δ-poly(L-histidine-cO-γ- benzyl ester of -L-glutamic acid)-b-poly(ethylene oxide)-b-poly(1-histidine-coγ- -L-glutamic acid benzyl ester)-b-poly(L-lysine), H2N-PEO-NH2 initiates copolymerization of monomers N<IM>-Trt-L-histidine N-carboxy anhydride (NCA, HIS-NCA) and u-L-glutamic acid gbenzyl ester NCA (BLG-NCA). All manipulations were performed under high vacuum in custom-made glass reactors, equipped with glass membranes, glass-coated magnets, and constrictions for the addition of reagents, following established high vacuum techniques.

[0074] The polymerization was carried out in a self-made apparatus, which has a 250 ml ground flask equipped with a high vacuum stopcock to periodically degas the solution and a glass-coated magnetic stirring bar. Also, the device has two self-made glass ampoules with two grinds for each of the ampoules for adding the monomers. The device is initially connected to the vacuum line through the burr and vented. The discharge of the device was tested using a Tesla machine and it was confirmed that the fabrication of the device using glass blowing techniques was successful. The device is then flame dried several times. It was then transferred to the argon chamber and 0.2 g of the bireactive acroreactive polyethylene oxide (NH2-PEO-NH2) which has a molecular weight = 20.0 x 10<3>g / mol (1.0 x 10<s>mol) was added. The device is connected to a vacuum line and degassed. Then 30 ml of ultra-dry benzene was distilled off and the macroinitiator was dissolved. The solution was stirred for two hours and the benzene distilled to dryness. The apparatus was left overnight in the high vacuum line. The next day 30 ml of high purity dimethylformamide (DMF) is distilled followed by dissolution of the macroinitiator. Then, the device was reinserted into the argon chamber, a mixture of 0.3384g of HIS-NCA (8.0 x10<-4>mol) and 0.2104g of BLG-NCA (8.0 x 10<-4>mol) was added to an ampoule equipped with grinding and artificial glass constriction. The device is connected to the vacuum line by grinding the ampoule containing the mixture, it was evacuated, followed by distillation of 5 ml of high-purity DMF, and the device was removed from the vacuum line by heat sealing the artificial glass constriction, and the mixture of the two monomers was dissolved . Addition of monomer was added by breaking the glass membrane of the ampoule, and the solution was stirred vigorously. The consumption of the monomers was monitored by FT-IR by removing a sample of the solution in the argon chamber. The solution is periodically pumped to remove the CO2 produced by the polymerization. Polymerization is carried out for 7 days. The monomer t-butyloxycarbonyl-L-lysine N-carboxy anhydride (Lys-NCA) was then added using the other ampoule and following exactly the same procedure. We added 3.27 x 10<-1>g of Lys-NCA (1.2 x10<-3>mol). Polymerization was continued for 3 days. After completion of polymerization, the polymer was precipitated in diethyl ether (300 ml) and filtered using a Buchner apparatus with a 0.45 pm filter. It is then dried under high vacuum (0.82 g 77% yield). The polymer was dissolved in CH 2 Cl 2 (5 ml) and then 10 ml of trifluoroacetic acid (TFA) was added. The polymer was allowed to deprotect for 1 hour at room temperature, and the solutions turned yellowish-light brown. Triethylsilane was then added dropwise until the solution turned from yellow-light brown to colorless. The solution was poured into diethyl ether (200 mL), and the white solid was filtered and dried. The white solid was then added to water and dialyzed against 2 L MilliQ water with p~5 (buffered with a dilute aqueous HCl solution) twice. It was then dialyzed against 2 liters of MilliQ water at pH ∼9 for four times and twice into pure MilliQ water. Finally, the polymer solution was freeze-dried to give the corresponding polymer as a white powder. The polymer finally obtained was 0.5 g (46% yield).

C. Results

C.1. Physicochemical Characterization of APTY.

The pentameric tetrapolymer was analytically characterized by a combination of methods including size exclusion chromatography, FTIR and <1>H NMR spectroscopy. The combined characterization methods revealed that the pentacomponent tetrapolymer was that expected from the stoichiometry and the synthesis design. APTY molecular characteristics are given in Table 1.

Table 1. Molecular Characteristics of the pentacomponent tetrapolymer PLys60-b-(PHis20-co-PBLG20)-PEO-b-(PHiS20-co-PBLG20)-b-PLys6

FOR. Rheological measurements to study hydrogel strength as a function of pH and temperature.

Rheology measurements were performed at two different pH values of 7.4 and 6.5, i.e. the pH of healthy and cancer tissue. At these pH values, rheological measurements were carried out at three different temperatures, i.e. 25, 37 and 40 °C. The results are shown in FIGs. 17-20. From the Dynamic Frequency Scan (DFS) plots it is evident that the hydrogel is responsive to different pH and temperature. The strength of the hydrogel is significantly affected by the temperature and pH between healthy and cancerous tissue. More importantly, hydrogel strength decreases going from healthy to cancerous tissue values. Consequently, this material is expected to respond even more targeted to cancerous tissues, and melt only towards the cancerous tissues compared to the healthy tissues, and finally, APTY will release targeted anticancer drugs selectively to the cancer tissues.

EXAMPLE 3

In Vivo and In Vitro studies of Hydrogem (hydrogel with entrapped gemcitabine).

A. Summary.

[0077] The in vitro study was inconclusive due to technical problems encountered under the specific experimental conditions. The in vivo study was performed using PTH hydrogel and hydrogem. Hydrojem consists of a hydrogel in which gemcitabine is entrapped. The amount of gemcitabine entrapped was 1:1 in relation to the weight of PTH ie the weight of the drug was equal to the weight of the polymer. The in vivo study was performed using human pancreatic cancer xenografts in immunocompromised muscles. Initial studies were performed using a hydrogel formed from poly(L-leucine) instead of poly(γ-benzyl-L-glutamate) in the PTH material. The results showed that the material with poly(L-leucine) caused toxicity at polymer concentrations above 100mg/kg. Acute toxicity studies on free PTH showed that the polymer material did not induce toxicity at concentrations up to 400mg/kg. Five different groups of muscles were used: a) untreated, i.e. they received neither polymer nor drug b) those that received the pure PTH polymer at the same concentration as hydrogem c) those that received gemcitabine intraperitoneally d) those that received gemcitabine subcutaneously peripheral to cancerous tumor and e) those who received hydrogem subcutaneously and peripheral to the tumor. The in vivo study showed that hydrojem is quite effective since it delays the growth of pancreas xenografts. This was seen as in the group of mice that received the hydrogem, at a concentration half that used for the free drug around the tumor, the reduction in tumor growth was comparable and even better. This indicates that the polymer selectively directs drug release to cancerous tissue rather than healthy tissue.

B. Methods.

A. In vitro studies. For the cytotoxicity study we used CHO (Chinese Hamster Ovarian cells) cells. On the first day, we placed 100μl of hydrogel in 9 wells of a 96-well cell culture microplate. The plate was then incubated at 37°C so that the hydrogel homogenized and covered the entire surface of the well. Then in three of the wells, we added 5000 CHO cells, in 100μl. CHO cells (5000 cells/well in 100μl culture medium, RPMI 1640, supplemented with antibiotics, glutamine and 10% fetal calf serum) were also cultured in the same dish as a control control for normal cell growth and/or for the addition of hydrogel after 24 cultivation hours. 24 hours after cell culture, we added 100μl of hydrogel to the wells where there were only cells. At the same time, in the wells that had only hydrogel, we added CHO cells at a concentration of 5000 cells/well. The culture continued for 24 hours and at the end we studied and compared the cell growth with that of cells cultured without the presence of hydrogel.

[0079] B. In vivo study. For the toxicity study in animal cancer models we used NOD-SCID (NOD.CB17-Prkdcscid/J) mice which we obtained from the Jackson Laboratory (The Jackson Laboratory 600 Main Street Bar Harbor, Maine 04609 USA). The colony was maintained in a pathogen-free environment in type IIL cages. Male and female muscles were used in the present study. All animals were maintained under specific pathogen-free conditions (SPF conditions) in our laboratory animal unit (EL42-BI002) with free access to water and food. The handling and experimentation of the animals was done in accordance with the current Greek legislation as well as the guidelines of the European Union and the European Council (86/609 and ETS123, respectively).

[0080] C. Acute toxicity study. For the acute toxicity study we followed the guideline from the National Cancer Institute (NCI, NIH, USA) for acute toxicity study protocol (http://dtp.nci.nih.gov/branches/btb/acute tox.html). One mouse receives an injection of 400mg/kg, a second mouse receives an injection of 200mg/kg and a third mouse receives an injection of 100mg/kg. Mice are observed for two weeks and sacrificed if they lose more than 20% of their body weight or if other signs of toxicity are observed. Mice have free access to food and water. The volumes of the substances under study for the required concentrations that had to be administered did not exceed 0.2 ml/ 10 g. animal weight. For the needs of the experimental approach, survival, body weight losses and changes in their behavior were studied and recorded for a period of two weeks.

[0081] D. Hydrogel toxicity study. In a preliminary study three female mice received a single concentration of hydrogel subcutaneously in the dorsal region while another received the same concentration intraperitoneally. Each of the mice received 200µl of the 1/40 hydrogel formulation in buffer. The muscles were monitored every day for two weeks. In a subsequent experiment one mouse received an injection of 200mg/kg (66µl of the 1/60 hydrogel formulation) and another received a single injection of 100mg/kg (330µl of the 1/60 hydrogel formulation). The concentration of 400mg/kg was not studied due to the large volume that would need to be obtained (1.2ml/mouse). As a negative control we used a mouse that received the solvent used to hydrate the hydrogel. The muscles were monitored every day for two weeks.

[0082] In a subsequent experiment we studied the accumulated toxicity from a hydrogel formulation with a concentration of 10mg/ml. For this experimental approach we used a total of two mice, one for each concentration and received a single injection of 200mg/kg (400μl of the hydrogel) and 100mg/kg (200μl of the hydrogel) respectively, intraperitoneally or subcutaneously. The muscles were monitored for two weeks.

[0083] Finally, in order to choose the most suitable form of hydrogel in which the drug gemcitabine will be encapsulated, we used eight muscles and divided them into two different groups (4 mice per group) and which received different concentrations of the two different forms of hydrogel containing poly (L-leucine) (PLEU) and poly(γ-benzyl ester of L-glutamic acid) (PBLG). The concentrations studied were 400, 200 and 100 (mg/kg). The injection of the hydrogel was done subcutaneously on the upper part of the back after we had shaved their fur. Each mouse from the two groups received a single injection of the respective form of hydrogel at a concentration of 400 mg/kg (320μl), 200 mg/kg (160μl), 100 mg/kg (80μl). The fourth mouse from each group received the corresponding form of hydrogel intraperitoneally at a concentration of 200mg/kg

[0084] E. Gemcitabine toxicity screening. In a first experimental approach we administered gemcitabine (Gemsar VL7501, Farmaserve-Lilly) at different concentrations to three male mice (one concentration/mouse). Thus 400, 200 and 100 (mg/kg) gemcitabine was administered to the muscles in a single injection, intrafascially or subcutaneously, and the muscles were monitored every day for two weeks, for possible signs of toxicity.

[0085] F. Sub-chronic toxicity testing of gemcitabine. The impending toxicity of gemcitapine was further studied and an appropriately designed efficacy study protocol was followed. For this purpose, three mice received two injections of gemcitabine in a period of one week, intrafascially or subcutaneously. Three different drug concentrations (concentration/mouse), 200, 100 and 50 (mg/kg) were studied.

[0086] G. Hydrogem activity study (hydrogel with gemcitabine)

The activity/efficacy of the gemcitabine hydrogel formulation (Hydrogem) was studied using human xenografts in mice. For this purpose we used the established human pancreatic cancer line AsPCl. For this experimental approach we used 35 NOD/SCID mice and each mouse was inoculated with 1x10<6> cells, subcutaneously, in the posterior axillary side. Mice were monitored every day until cancerous tumors developed. During the appearance of the cancerous tumors and with the use of the calliper we measured and recorded the two diameters, twice a week. Using the formula (axb2/2) where a= larger and b= smaller, diameter, the size of each tumor was calculated. A graph of tumor size versus time (in days after cell embolization) was plotted to study tumor growth. When the size of the tumors reached about 200mm<3> the mice were randomly divided into five different groups, where each group included seven muscles (7 tumors/group). One of the groups received nothing and was used as a reference group. Two subsequent groups received gemcitabine, subcutaneously or intraperitoneally respectively, at a point close to the area of the cancerous tumor (peritumoral, "pt") where they served as positive controls. The fourth group received hydrogel while the fifth received hydrogem, both peripheral to the tumor. The muscles in the two gemcitabine groups received the chemotherapy drug twice with an interval of one week at a concentration of 100mg/kg. The hydrogem group received the formulation in a single administration (83μl), an amount corresponding to the administered concentration of gemcitabine while the muscles in the hydrogel group also received 83μl of the formulation. Tumor growth as well as muscle survival were monitored and the corresponding data recorded. When the size of the tumors exceeded 10% of the body weight of the respective mouse then the animal was killed using ether and the experiment was terminated when all the muscles were killed. As an indication of toxicity, we measured the weights of the animals, twice a week, as well as observed any other signs that could indicate a similar consequence.

C. Results.

[0087]      The use of the in vitro technique, for the study of imminent toxicity, did not produce any result regarding the activity of the hydrogel due to technical problems. The hydrogel did not allow the cells to attach successfully to the culture dishes, resulting in their death. Because of this we did not continue the in vitro toxicity study.

[0088]         The in vivo toxicity study indicated, as the most suitable form of hydrogel to encapsulate the drug, PBLG at a concentration of 1/40. PLEU 1/40 was toxic as the mouse that received it intraperitoneally died two days after the injection. In addition, muscles receiving 400 and 200 (mg/kg) PLEU showed extensive skin irritation at the injection site, and PLEU disappeared shortly after injection indicating the instability of the formulation (FIG. 21). In contrast to PLEU, PBLG showed more favorable properties. It showed no toxicity at any of the concentrations studied, while a minor skin irritation at the injection site appeared to be the most obvious side effect. In addition, PBLG remained visible at the site of administration for a long time indicating its suitability as an entrapment agent for gemcitabine (FIG. 22).

[0089] The subsequent activity study indicated the combination of hydrogel with gemcitabine more effective than the free drug for both different routes of administration (intraperitoneal and subcutaneous - ip and sc respectively-), as it was seen under the specific experimental conditions. Indeed, the administration of gemcitabine encapsulated in the polymer, at half the dose of the free form of the drug, could delay the growth of cancerous tumors to the same extent as shown by the growth curves (FIGS. 23 and 24). Hydrogem showed a comparable and even better %ΔT/ΔC value compared to free gemcitabine (ip or sc), while the tumor doubling time increased correspondingly to the tumor doubling time in the muscle group treated with free gemcitabine, regionally of the volume, for two times with an interval of one week. Finally, through the Kaplan-Meier survival curve, it appeared that hydrogem was more suitable than free gemcitabine as the survival time of the muscles that received it increased more than the group of muscles that received free gemcitabine (FIG.25). In addition, the hydrogem remained attached to the cancerous tissue rather than to the healthy one, thus confirming the selective transport of the drug.

Claims (20)

Claims: 1. A linear pentameric polymer of the formula A-b-B-b-C-b-B-b-A, consisting of at least one polypeptide block. Clusters B are hydrophobic while clusters A and C are hydrophilic. Hydrophobic clusters are composed of pH-responsive polypeptides and contain a significant proportion with α-helix secondary structure and a smaller proportion of random coil and β-sheet structure. The pentapolymer is capable of self-healing by converting from the liquid phase to a hydrogel in less than 10 seconds when the application of shear stress is stopped, and is pH responsive and able to convert from the hydrogel state to the liquid state under the influence of different pH values in the environment of. 2.   A linear pentameric polymer with the basic structure and properties according to claim 1, which is of the type poly(E)-b-poly(L-histidine-co-Δ)-b-poly(E)-b- poly(L-histidine-co-Δ)-b-poly(E). 3.   A linear block polymer according to claim 2, wherein the poly(E) moiety is poly(L-lysine) and D is from one or more polypeptides selected from a group consisting of poly(γ-benzyl ester of L -glutamic acid), poly(L-alanine), poly(L-glycine), poly(L-valine), poly(L-phenylalanine), poly(L-isoleucine), poly(L-methionine), poly(L -tryptophan), and poly(L-β-benzyl ester of aspartic acid). 4. A linear multicomponent polymer according to claim 3, selected from one of the following structures: poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine), poly(L-lysine)-b-poly(L-histidine-co-L-alanine)-b-poly(L-lysine)-b-poly( L-histidine-co-L-alanine )-b-poly ( L-lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-glycine)-b-poly(L-lysine)-b-poly(L-histidine-co-L-glycine )-b-poly ( L-lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-valine)-b-poly(L-lysine)-b-poly(L-histidine-co-L-valine)-b-poly( L-lysine), poly(L-lysine)-b-poly(L-histidine-co-L-phenylalanine)-b-poly(L-lysine)-b-poly(L-histidine-co-L-phenylalanine )-b-poly( L-lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-isoleucine)-b-poly(L-lysine)-b-poly(L-histidine-co-L-isoleucine)-b-poly( L-lysine), poly(L-lysine)-b-poly(L-histidine-co-L-methionine)-b-poly(L-lysine)-b-poly(L-histidine-co-L-methionine )-b-poly ( L-lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-tryptophan)-b-poly(L-lysine)-b-poly(L-histidine-co-L-tryptophan )-b-poly { L-lysine), poly(L-lysine)-b-poly(L-histidine-co-γ-aspartic acid benzyl ester)-b-poly(L-lysine)-b-poly(L-histidine-co-γ-aspartic acid benzyl ester )-b-poly(L-lysine). 5.   A linear multicomponent polymer according to claims 1 or 2 wherein the block A, C and poly(E) may be one or more polypeptides selected from a group consisting of poly(L-serine), poly(L- arginine), poly(L-threonine), poly(L-asparagine), poly(L-glutamine), and D is γ-benzyl ester of L-glutamic acid. 6. A linear block polymer according to any of claims 1, 2 and 5 selected from one of the following structures: poly(L-serine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-serine), poly(L-arginine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine) -b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-arginine), poly(L-threonine )-b-poly( L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-lysine )-b-poly ( L-histidine-co-γ-benzyl ester of L-glutamic acid)-b-poly(L-threonine), poly(L-asparagine )-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly( L-lysine )-b-poly( L-histidine-co-γ-benzyl ester of L-glutamic acid)-b -poly(L-asparagine), poly(L-glutamine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ether of L- glutamic acid)-b-poly(L-glutamine), 7. A linear block polymer according to any one of claims 1 to 3, which is poly(L-lysine)(30-90)-b-poly(L-histidine(10-30)-co-γ-benzyl ester of L-glutamic acid(10-30))-b-poly ( L-lysine )(100-300)-b-poly ( L-histidine(10-30)-co-γ- benzyl ester of L-glutamic acid ( 10-30))-b-poly(L-lysine)(30-90). 8. A linear block polymer of the formula A-b-B-b-C-b-B-b-A according to claim 1, whose block C is poly(ethylene oxide) and consists of at least one polypeptide block. Clusters B are hydrophobic while clusters A and C are hydrophilic. Hydrophobic clusters are composed of pH-responsive polypeptides and contain a significant proportion with α-helix secondary structure and a smaller proportion of random coil and β-sheet structure. The pentablock polymer is capable of self-healing by converting from the liquid phase to a hydrogel in less than 10 seconds when the application of shear stress is stopped, and the block polymer is pH responsive and capable of converting from the hydrogel state to the liquid state under the influence of different pH values in its environment. This polymer exhibits, in addition to the characteristics and properties mentioned in claim 1, the property of being thermoresponsive and able to transform from the hydrogel state to a liquid state under different conditions of temperature and/or pH change. 9. A linear block pentamer with the basic structure and properties according to claim 8, which is of the type poly(E)-b-poly(L-histidine-co-Δ)-dpoly(ethylene oxide)-b- poly(L-histidine-co-Δ)-b-poly(E), wherein the hydrophobic portion D is one or more polypeptides selected from a group consisting of poly(γbenzyl ester of L-glutamic acid), poly(L-alanine ), poly(L-glycine), poly(L-valine). poly(L-phenylalanine), poly(L-isoleucine), poly(L-methionine), poly(L-tryptophan), and poly(L-β-benzyl ester of aspartic acid). 10. A linear block polymer according to claims 8 and 9, the A and poly(E) segment is poly(L-lysine), and is selected from one of the following structures: poly(L-lysine)-b-poly(L-histidine-co-L-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-L-γ-benzyl ester of L-glutamic acid)-bpoly(L-lysine), poly(L-lysine)-b-poly(L-histidine-co-L-alanine)-b-poly(ethylene oxide)-b-poly(L-histidine-co-L-alanine )-b-poly ( L- lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-glycine)-b-poly(ethylene oxide)-b-poly(L-histidine-co-L-glycine)-d-poly (L- lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-valine)-b-poly(ethylene oxide)-b-poly(L-histidine-co-L-valine )-b-poly ( L- lysine ), poly(L-lysine)-b-poly(L-histidine-co-L-phenylalanine)-b-poly(ethylene oxide)-b-poly(L-histidine-co-L-phenylalanine)-b-poly(L- lysine), poly(L-lysine)-b-poly(L-histidine-co-L-isoleucine)-b-poly(ethylene oxide)-b-poly(L-histidineco-L-isoleucine)-b-poly(L-lysine) , poly(L-lysine)-b-poly(L-histidine-co-L-methionine)-b-poly(ethylene oxide)-b-poly(L-histidineco-L-methionine)-b-poly(L-lysine) , poly(L-lysine)-δ-poly(L-histidine-co-L-tryptophan)-b-poly(ethylene oxide)-b-poly(L-histidine-co-L-tryptophan)-b-poly(L- lysine), poly(L-lysine)-b-poly(L-histidine-co-γ-aspartic acid benzyl ester)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-aspartic acid benzyl ester)-b -poly(L-lysine). 11. A linear polyblock polymer according to claims 8 or 9 wherein the blocks A and poly(E) are from one or more polypeptides selected from a group consisting of poly(L-serine), poly(L-arginine), poly(L-threonine), poly(L-asparagine), poly(L-glutamine). A linear block polymer according to any one of claims 8, 9 and 11, which is selected from one of the following structures: poly(L-serine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid )-bpoly(L-serine), poly(L-arginine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid s)-bpoly(L-arginine), poly(L-threonine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid )-bpoly(L-threonine), poly(L-asparagine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid )-bpoly(L-asparagine), poly(L-glutamine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(ethylene oxide)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid )-bpoly(L-glutamine), 13. A linear block polymer according to claim 8, which is poly(L-lysine)-b-poly(L-histidine-co-γ-benzyl ester of L-glutamic acid s)-bpoly(ethylene oxide)-b -poly(L-histidine-co-γ-benzyl ester of L-glutamic acid)-bpoly(L-lysine). 14. A linear block polymer according to any one of claims 5 to 7 which is poly(L-lysine)(30-90)-b-poly(L-histidine(10-30)-co-γ-benzyl ester of L-glutamic acid(10-30))-b-poly(ethylene oxide)(200-700)-b-poly(L-histidine(10-30)-co-γbenzyl ester of L-glutamic acid(10-30)) -b-poly(L-lysine)(30-90). 15.  Use of the linear blocky multicomponent polymer described in claims 1 to 7 in drug delivery applications. 16.  A hydrogel consisting of linear block polymers according to claims 1 to 15 and at least one pharmaceutical compound. 17. A hydrogel according to claim 16, wherein the pharmaceutical compound is selected from a group of substances consisting of gemcitabine, paclitaxel, doxorubicin, everolimus, cisplatin, epirubicin. , daunorubicin, topotecan, fluorouracil, camptothecin, si-RNA, and plasmid DNA. 18. A hydrogel according to claim 16 or 17, wherein the pharmaceutical compound is gemcitabine. 19.  A hydrogel according to claims 16 or 17 for use in the fight against cancer. 20. A hydrogel according to claims 16 to 18 for use in combating pancreatic cancer.