membranes for transdermal delivery of therapeutics

Polysaccharide hydrogel films/ membranes for transdermal delivery of therapeutics

22

Zahra Shariatinia, Azam Barzegari Department of Chemistry, Amirkabir University of Technology (Tehran Polytechnic), Tehran, Iran

22.1 Introduction Skin has three layers including the epidermis which is the outermost layer of skin providing a waterproof barrier and which creates our skin tone [1]. The dermis is placed below the epidermis and contains tough connective tissue, hair follicles and sweat glands. The deeper subcutaneous tissue (hypodermis) is made up of fat and connective tissue. The skin usually interfaces with the environment, thus it plays a significant immunity role to protect the body against pathogens and excessive water loss [2]. Moreover, skin can function as an insulator, sensing system, and temperature regulator. It is also a place for the synthesis of vitamin D and protection of vitamin B folates. Very damaged skin tries to heal through formation of scar tissue which is usually discolored and depigmented [3]. The skin is the most easily accessible tissue of the body which provides a barrier for the micro and macromolecules existing in the environment due to its low permeability to these compounds [4]. The skin surface area of an average adult human body is about 2 m2, and it contains around one-third of the total blood circulating in the body. Percutaneous absorption of drug by the skin mostly happens through stratum corneum which is made up of dead and keratinized epidermal cells with a thickness of 10 μm which can act as a barrier against penetration of drug molecules. Consequently, the passage of drugs across the skin is difficult [5]. Localized on-demand drug delivery systems can exactly control the drug release rate [6,7] and they have promising application in personalized medicine to match an individual patient with the most helpful medical treatments [8]. Drug administration through the skin is performed in order to treat skin diseases topically or for the transdermal absorption of drugs into the systemic circulation. The topical route provides a large and diverse surface along with the comfort in application by self-administration and offers a substituting method to both hypodermic injection and oral drug delivery [9]. The amount and rate of drug absorption from the skin is dependent to the skin physiology, and the physicochemical features of both drug and the delivery system. The present dosage forms such as patches, ointments, and creams have some limitations. Patches have numerous drawbacks including the usual skin irritation [10], having occlusive properties leading to obstruction of sweat ducts thus preventing loss of water vapor from the skin surface, Polysaccharide Carriers for Drug Delivery. https://doi.org/10.1016/B978-0-08-102553-6.00022-2 Copyright © 2019 Elsevier Ltd. All rights reserved.

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problematic application on the curved surfaces, pain when they are peeled off and undesirable aesthetic appeal [11]. The creams and ointments, which are semisolid formulations, overcome some of these such disadvantages, but they have other limitations. They do not have permanent contact with the skin surface and can simply be wiped off by the patient’s clothes. Therefore, repeated utilization is necessary in the case of chronic diseases including athlete’s foot, candidiasis, and ringworm. Furthermore, they leave a greasy and sticky feel after usage which results in poor patient satisfaction [12]. Thus transdermal delivery systems and permeation improvement methods are widely investigated to tackle these barriers by improving the drug transdermal delivery [13]. Hence, it is required to develop a dosage form which certificates less common dosing through preserving close contact with the skin over a long time in order to increase the patient compliance. The film forming system is a new method employed as a substitute for the common transdermal and topical formulations. These systems are nonsolid dosage forms that produce films in situ after they are applied onto the skin or any surface on the body. Such a system is composed of drug and film forming excipients in a vehicle so that when it contacts the skin it can leave a film of excipients together with the drug after solvent evaporation. Indeed, the film forming method is directly applied to the skin to form a thin, transparent film in situ by solvent vaporization. The formed film is either a solid polymeric material which can act as a matrix for the sustained drug release into the skin or it is a residual liquid film that is quickly absorbed into the stratum corneum [14]. After loss of the volatile components of the vehicle formulation and film formation on the skin surface, the drug concentration enhances and reaches saturation level with the probability of attainment to the supersaturation level on the skin surface. This can increase the drug flux through the skin by improving the thermodynamic activity of the vehicle without influencing the skin’s barrier role, thus decreasing the side effects and irritation as well as overcoming the problem of instability [15]. Accordingly, this film increases drug penetration through skin relative to other transdermal dosage forms confirming the film forming system is an intermediate between the transdermal patches and semisolid dosage forms which displays the benefits of both systems. Film forming systems were mostly used for surgery and wound care purposes, such as tissue glues for the closure of operative wounds, which can contain drugs and antimicrobial agents to inhibit infections in the wounds [16]. The film forming materials can be natural (such as fibrin) or synthetic (such as cyanoacrylates) and they may be applied in nonmedical uses including delivery of active ingredients added into beauty products such as silicone film forming materials used to make cosmetic ointments and creams. Also, the film forming materials are utilized as transparent peel off masks to treat skin hydration and acne problems. Another application of the film forming agents is that they can act as substrates (e.g., hydrophobic and hydrophilic ointments and creams as well as UV protecting creams) for numerous barrier membranes in industry to protect workers from acids, bases, detergents, various hazardous chemicals, UV exposure, and infra-red heat [17]. Furthermore, the film forming polymers can be sprayed onto the soil to form membrane films in order to increase the soil integrity and to elevate the soil temperature, which is beneficial in crop protection [18]. Polysaccharides are biocompatible natural polymeric compounds having modifiable functional groups on their chains which make them perfect candidates for

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biomedical applications [19]. Their film forming ability, biodegradability, biocompatibility, and nontoxicity have led to substantial attention. Hydrogels are polymers which are swollen in water and possess three-dimensional network structures composed of hydrophilic polymeric chains. They have hydrodynamic properties similar to those of biological tissues which can be used as drug delivery systems since hydrogels excellently disperse drugs inside the polymeric matrix [20]. Hydrogels can be used in different forms such as creams and patches for oral administration, injection, and transdermal delivery [21]. They assist the drug permeation into the skin by skin hydration with a moisturizing influence indicating they are appropriate for topical applications. Additionally, they can act as stabilizing agents to improve the transdermal delivery of several systems such as micelles, liposomes, and nanoparticles [22,23]. Hydrogels prepared from a natural polymer such as chitosan, hydroxyethylcellulose, carboxymethyl cellulose, gelatin, gellan gum, guar gum, acacia gum, pectin, hyaluronic acid, chondroitin sulfate, collagen, and alginate have improved availability in the biomedical arena as extremely biocompatible biomaterials [24]. Smart hydrogels can serve as potential carriers for on-demand drug delivery that displays stimulus-responsive behaviors [25], including pH [26], temperature [27], mechanical force [28–30], magnetic fields [31,32], and light [32–35]. It is noteworthy that numerous efforts are being carried out to achieve smart hydrogels using degradable natural biomaterials to be employed in localized on-demand drug delivery. In this chapter, polysaccharide hydrogel films/membranes are described as promising candidates for topical and for transdermal delivery of therapeutics. For this purpose, various types of such systems are reviewed including chitosan, hydroxyethylcellulose, carboxymethyl cellulose, gelatin, gellan gum, guar gum, acacia gum, pectin, hyaluronic acid, chondroitin sulfate, collagen, alginate, and smart hydrogels along with their characteristics and applications. Furthermore, the results of our research on the fabrication of electrospun nanofibrous three-layered chitosan-polyethylene oxide/ polylacticacid/chitosan-polyethylene oxide (CS-PEO/PLA/CS-PEO) mats containing green tea (GT) are presented and discussed.

22.2 Chitosan drug delivery systems Injectable hydrogels are biodegradable and in situ formable which can effectively and homogeneously encapsulate drugs/cells in a minimally invasive way. Up to now, several biodegradable, biocompatible, and injectable hydrogels have been produced through pH, temperature, chemical, and UV-irradiation triggered gelation methods for localized drug delivery that have shown minimal invasion with different release behaviors for the therapeutic drugs [36]. Chitosan (CS) is a natural derivative of chitin and a polysaccharide composed of copolymers of glucosamine and N-acetylglucosamine units that are linked via β-(1,4)-glucosidic bonds. It is a weak base having an inherent pKa ~ 6.3 and a low charge density. It is a biodegradable, natural biomaterial containing amino groups, which is extensively applied in tissue engineering, localized drug delivery, and injectable hydrogels for cancer therapy because it has valuable biological features such as biocompatibility, biodegradability by lysozyme, hemostatic ability,

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and antibacterial activity. Injectable CS is undergone to thermal and pH triggered gelation and could be degraded in vivo by means of chitosanase and lysozyme enzymes. Also, due to poor solubility of native CS in the neutral pH medium, its applications are limited and consequently it fails to achieve cell-laden hydrogel at physiological pH. Thus in order to address this subject, recently UV cross-linking ability was incorporated into CS to allow fabrication of patterned cell-laden and fast in vivo transdermal curing hydrogel. Hydrosoluble, UV cross-linkable, and injectable N-methacryloyl chitosan (N-MAC) was synthesized and led to fabricate cell-laden microgels with on-demand patterns by means of photolithography and it was exhibited that low-dose UV irradiation scarcely prompted skin injury and severe inflammatory response was vanished after 7 days [37]. In another work, CS/hyaluronan transdermal films were developed using aqueous solutions containing diverse weight ratios of hyaluronan and CS to improve bioavailability of thiocolchicoside and it was found that when the hydration of the polymeric network was lower, the drug diffusion inside the films and its penetration through the skin were lower [38]. Rabeprazole-alginate core coated CS nanoparticles (NPs) were produced using water in oil nanoemulsion method and then transdermal rabeprazole-NPs loaded patches were developed to prevent drug peroral acid sensitivity and first pass influence [39]. In vitro permeation of the NPs was evaluated in comparison with rabeprazole. Also, ex vivo permeation of patches into rat skin was investigated by studying the permeation mechanism and kinetic analysis. The optimized rabeprazole-NPs displayed a sustained release behavior. Patches loaded with rabeprazole-NPs revealed high skin permeability as well as controlled drug release. Some nanocomposite films made up of CS, starch, glycerin, cyclophosphamide drug, and Fe3O4 NPs (thickness 0.13–0.2 mm) were fabricated containing several amounts of Fe3O4 NPs (2%–10%) [40]. The in vitro bactericidal activities against Gram-negative Escherichia coli and Gram-positive Staphylococcus aureus bacteria were done and greater antibacterial effects were achieved against Gram-positive bacterium. Biocompatible nanocomposite films of CS and poly(ethylene glycol)-block-poly(propylene glycol)-block-poly(ethylene glycol) (BP) polymers composed of metformin (MET) drug and MCM-41 or MCM-41-APS (APS = aminopropylsilane) NPs were fabricated as drug delivery systems which were beneficial for controlled drug release [41]. The MET release was sharply enhanced during ~22–24 h (burst release) but after that the drug release was slowly increased within 15 days (sustained release). In another work, antimicrobial chitosan–polyethylene oxide (CS-PEO) nanofibrous mats containing 3%, 5%, and 10% (w/w) of zeolitic imidazolate framework-8 (ZIF-8) NPs (diameter ∼60 nm) were produced by the electrospinning method [42]. The CSPEO-GA-3% ZIF-8 NPs cross-linked with glutaraldehyde (GA) was also fabricated. Antimicrobial activities of the CS-PEO and CS-PEO-3%ZIF-8 mats were tested using the viable cell-counting method to determine their efficiency in dropping or stopping the growth of E. coli and Staphylococcus aureus bacteria and it was found that the CS-PEO mat comprising 3%ZIF-8 had 100% bactericidal activity against both kinds of microorganisms. Antibacterial CS-PEO nanofibrous mats were fabricated through the electrospinning method for wound dressing applications using a green route for the development of mats which were loaded with 0.25% and 0.50% (w/w) of bioactive

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silver (Ag) NPs (diameter ∼70 nm) reduced by Falcaria vulgaris herbal extract [43]. Both of the CS-PEO mats containing 0.25% and 0.50% bioactive F. vulgaris-Ag NPs exhibited 100% bactericidal effects against both E. coli and Staphylococcus aureus bacteria. The release of Ag NPs from both CS-PEO mats including 0.25% and 0.50% F. vulgaris-Ag nanofibrous mats was suddenly enhanced during the first 8 h while after that the Ag NPs were released very sluggishly (practically constant). To design effective wound dressing materials, antimicrobial electrospun CS-PEO nanofibrous mats incorporated with ZnO NPs and hydrocortisone-imipenem/cilastatinloaded ZnO NPs were fabricated [44]. The hydrocortisone release was enhanced to 82% during 12 h but the release rate of imipenem/cilastatin was so much slower (20% of the drug was released during 12 h) signifying this nanofibrous mat is very appropriate for the inhibition of both infection (using imipenem/cilastatin antibiotic and ZnO NPs) and inflammation (by hydrocortisone) predominantly for wound dressing application. Recently, antimicrobial electrospun CS-PEO nanofibrous mats incorporated with cefazolin, fumed silica (F. silica) and cefazolin-loaded F. silica NPs were created to be used for wound dressing application [45]. The CS-PEO mats comprising 2.5% cefazolin and 1% F. silica-0.50% cefazolin displayed 100% antibacterial effects against both E. coli and Staphylococcus aureus bacteria. The cefazolin release from mats was abruptly increased in 24 and 6 h from the CS-PEO mats containing 2.5% cefazolin and 1% F. silica-0.50% cefazolin and then the drug was released very sluggishly. The ability of the CS-PEO-F. silica-cefazolin mat to heal wounds as a wound dressing was assessed on the wounded skins of the female Wistar rats and it was revealed that the wounded skins of the rats were practically completely healed after 10 days by means of this efficient mat. An earlier study aimed to improve the systemic bioavailability of propranolol-HCl by designing a transdermal drug delivery system based on freeze dried CS NPs dispersed into gels made of poloxamer and carbopol [46]. The CS NPs were achieved through ionic gelation method using tripolyphosphate cross-linker. It was indicated that the smallest propranolol loaded CS NPs were obtained using 0.2% CS and 0.05% tripolyphosphate. The CS NPs dispersed into the poloxamer-carbopol gel displayed a thixotropic property with a lengthy drug release behavior which was confirmed through the permeation tests into pig ear skin. Some antimicrobial nanocomposite films of CS/phosphoramide/1%–5% Fe3O4 NPs were produced for application using in biomedical areas [47]. The phosphoramide was synthesized by the reaction of morpholine with N-4-nitrophenyl phosphoramidic dichloride. The in  vitro bactericidal effects were assessed against four bacteria: two Gram-positive Staphylococcus aureus, Bacillus cereus and two Gram-negative E. coli, Pseudomonas aeruginosa microorganisms and superior antibacterial activities for the films against Gram-positive bacteria. In order to prepare platforms for transdermal drug delivery, CS polysaccharide was mixed with phospholipids to produce electrospun hybrid nanofibers [48]. The nanofibers were stable in the PBS solution for at least 7 days. Fluorescence microscopy proved that L929 cells seeded on the surface of the CS/phospholipids hybrid had similar metabolic activity which was comparable to the cells seeded on a tissue culture plate (control). The release of diclofenac, curcumin and vitamin B12 (as model drugs) from the CS/phospholipids hybrid nanofibers was performed and it was found that they were promising transdermal drug delivery systems.

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CS was modified through amide bond formation with sulfhydryl compound thioglycolic acid. Afterwards, preactivated CS conjugate was recognized by preactivation of CS thioglycolic acid with mercapto nicotinamide which was covalently linked via the disulfide bond. All conjugates were tested for their dermal adhesiveness and controlled drug release features. It was found that the preactivated CS conjugates displayed 7.46-fold dermal adhesiveness on skin due to their high tensile adhesion strengths. Additionally, 1.9-fold controlled release of Rhodamine123 drug was measured compared with unmodified CS [49]. It is known that wound healing is a dynamic process wherein cells and macromolecules work together in order to expedite tissue regeneration and repair tissue integrity and healing of the full thickness (FT) wounds necessitates extra support from native or synthetic matrices to assistance tissue regeneration, especially a matrix with optimum hydrophilic-hydrophobic balance which can swell adequately and decrease bacterial adhesion [50]. Hence, in a research effort, polyurethane diol dispersion and the antibacterial CS were blended in diverse ratios and self-organized to create macroporous hydrogel scaffolds (MHS) at room temperature after drying. In vitro cytocompatibility was verified through the proliferation of primary rat fibroblast cells on MHS. Also, MHS was used for an in vivo FT wound healing study on Wistar rats and it was compared with a similar commercial polyurethane containing dressing (Tegaderm). The MHS-treated wounds revealed faster healing along with enhanced wound contraction, greater collagen synthesis, and vascularization in the wound area compared to Tegaderm [51]. In another effort, CS/phosphoramide (pH)/Ag NPs nanocomposite films were developed comprising 1%–5% Ag NPs [52]. The in vitro antibacterial tests were assessed against four bacteria including two Gram-positive B. cereus, Staphylococcus aureus and two Gram-negative Pseudomonas aeruginosa, E. coli microorganisms and greater antibacterial effects were measured for the films against Gram-positive bacteria.

22.2.1 Chitosan-based hydrogels for wound healing applications Wound healing is a complex biological process which is related to an integrate response of different cell types and growth factors [53]. In order to accelerate the wound healing, biomaterial dressings are frequently employed [54]. In fact, an ideal wound dressing should maintain hemostasis, moist environment at the wound, prevent diffusion of dust and bacteria and remove excess exudates. Moreover, it should permit water and air permeation and promote epithelization by releasing biological agents to the wound [55]. It should also be nontoxic, nonallergenic, nonadherent, and easily removed without trauma. Also, it should be prepared from a readily available biomaterial which requires minimal processing, possesses antimicrobial properties and promotes wound healing [56]. Bacterial infection is one of the main reasons which complicates and postpones the wound healing process. Therefore, there is an increasing demand for advanced antimicrobial wound dressings within the wound care market. Such dressings contain antimicrobial components which are released to the wounded area to permit antibacterial activity by maintaining a healthy concentration to the healing tissues [57]. Current investigations emphasize the acceleration of the wound repair using systematically designed dressing materials, particularly biologically derived materials such as chitin, chitosan, and their derivatives, which are capable of accelerating the healing processes

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at molecular, cellular, and systemic levels. Antimicrobial herbal extracts are widely used in wound treatment without damaging the healthy tissues [58]. Green tea (GT) is one of the most extensively consumed beverages in the world due to its valuable health properties [51]. The major constituents of the GT leaves are polyphenols and the most favorable effects of GT are ascribed to them, particularly the catechins, which make up 25%–35% of the dry weight of GT leaves [59]. Catechins exist in plants and in foods of plant origin including beverages including GT which contains a considerable amount of catechins (Fig. 22.1A), namely, epicatechin (EC), epigallocatechin (EGC), and their gallate esters: epicatechin gallate (ECG), and epigallocatechin gallate (EGCG), among which the latter is the most abundant. These polyphenolic compounds contribute significantly to the beneficial health effects of GT.

22.2.1.1 CS-PEO/PLA/CS-PEO nanofibrous mats containing GT extract SEM micrographs and XRD patterns Herein, CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT extract were fabricated by the electrospinning method (Fig.  22.1B). The digital photograph of the instrument is shown in Fig. 22.1C. The SEM images presented in Fig. 22.2A–E belong to CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%, 1%, 2%, 3%, 4% (w/w) of GT, respectively. It is observed in all cases that very uniform fibers free of any nanobeads and droplets are produced, and their diameters are between 80 and 90 nm. It should be noted that in order to achieve such nanofibrous mats a series of experiments must be performed to find the optimum electrospinning conditions including applied voltage and the distance of the syringe from the drum collector. The XRD patterns of CS, PEO, PLA polymers, and CS-PEO-4% GT/PLA/CSPEO-4% GT nanofibrous mat are given in Fig. 22.3. The CS powder shows two sharp peaks at 10.48 degrees (25% intensity) and 19.93 degrees (100% intensity) suggesting the formation of inter- and intra-molecular hydrogen bonds in the presence of free amino and hydroxyl groups in CS. Also, a broad peak with low intensity is observed between 35 and 45 degrees which can be attributed to the rather amorphous nature of CS. It is notable that the crystallinity of CS is due to the formation of hydrogen bonds between its amino, hydroxyl, and N-acetyl groups leading to the appearance of sharp peaks in the X-ray crystallographic pattern which are generated from parallel and antiparallel alignments of polymeric chains or sheets. The PEO powder displays a crystalline XRD pattern containing two sharp peaks at 19.36 degrees (90% intensity), 23.52 degrees (100% intensity) plus two adjacent peaks with low intensities at 26.12, 26.39 degrees and another two peaks with very low intensities at 35.59 and 39.61 degrees. The PLA powder reveals a broad peak with the highest intensity at 16.58 degrees. The XRD pattern of the CS-PEO-4% GT/PLA/ CS-PEO-4% GT has a broad peak with the highest intensity at 16.67 degrees.

Tensile strength The tensile strength diagrams of CS-PEO/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT are presented in Fig. 22.4A. In fact, the tensile strength analysis was done in order to explore how much stress the samples will bear before

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Polysaccharide Carriers for Drug Delivery

OH OH HO

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(C) Fig. 22.1  (A) The molecular structures of main catechins exist in the green tea extract. Digital photographs of (B) the electrospun CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%, 1%, 2%, 3%, and 4% of GT extract and (C) the electrospinning instrument used in this study.

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Fig. 22.2  The SEM images of CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%, 1%, 2%, 3%, and 4% of GT extract.

suffering permanent deformation/tearing. It is evident that CS is a rigid and brittle natural polymer but PEO and PLA chains are flexible, thus it is expected that the CS-PEO/PLA/CS-PEO mats with a three-layered structure will show increased flexibility. Furthermore, the effect of GT extract addition to these mats is evaluated on the strength properties. It is seen in Fig. 22.4A that the highest elongation at break (maximum tensile strain on horizontal axis) is obtained for the mat loaded with 2% GT (6.17%) indicating the

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PEO Intensity (a.u.)

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Fig. 22.3  The XRD patterns of CS, PEO, PLA polymers, and CS-PEO-4% GT/PLA/CSPEO-4% GT nanofibrous mat.

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Fig. 22.4  (A) The tensile strength diagrams of CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%, 1%, 2%, 3%, and 4% of GT. The (B) TGA and (C) DSC diagrams of PLA polymer and CS-PEO, CS-PEO/PLA/CS-PEO and CS-PEO-4% GT/PLA/CS-PEO-4% GT nanofibrous mats.

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highest flexibility of the mat. On the contrary, the least elongation at break belongs to the mat containing 4% GT (4.25%) which can be related to the greatest amount of GT extract preventing very high flexibility for this sample. Nevertheless, it must be noted that these data are only comparative and in practice all of the nanofibrous mats containing 0%–4% of the GT herbal extract exhibit high flexibilities and tensile strengths with no fracturing even upon 180 degrees inward/outward bendings. Therefore, this result merely confirms that the flexibility is increased by addition of 1%–3% GT to the CS-PEO blend in the CS-PEO-GT/PLA/CS-PEO-GT mats but further GT addition (4%) results in decreasing the elongation at break. The possible reason for this behavior may be related to the highly enhanced inter- and intra-molecular hydrogen bonding interactions of CS and PEO functional groups with those of GT extract (catechins). The increased interactions occur between amino (NH2), acetyl (CH3CO), and hydroxyl (OH) groups of CS as well as OH, CO moieties of PEO chains with the OH, CO and CO functional groups of catechins in the GT which lead in decreasing the tensile strain for the mat containing 4% GT. Fig. 22.4A also illustrates that the tensile strength (the maximum value of vertical axis) of the CS-PEO-GT/PLA/CS-PEO-GT mat is decreased from 18.39 MPa (in the mat without GT) to 10.52 MPa (in the mat loaded by 1% GT) but further addition of GT (2%–4%) increases in the tensile strength so that it is the highest for the mat having 4% GT (29.64 MPa) while it is the lowest for the mat with 1% GT. The Young’s modulus, which is the slope of the graphs, has a similar increasing trend to that of tensile strength with the highest Young’s modulus measured for the mat including 4% GT. As a result, it can be found that 3% content of GT extract produces a mat with appropriate flexibility and high tensile strength. According to the tensile test results, the CS-PEO-3% GT/PLA/CS-PEO-3% GT is selected as the best mat among all of the samples examined in this work.

TGA/DSC analysis The thermal stabilities of the PLA polymer and CS-PEO, CS-PEO/PLA/CS-PEO, and CS-PEO-4% GT/PLA/CS-PEO-4% GT nanofibrous mats were examined by the TGA/DSC analysis and the corresponding diagrams are given in Fig. 22.4B. It is obvious from the TGA graphs that the PLA degradation starts at approximately 300°C and terminates at about 375°C indicating 100% weight loss percent by thermal decomposition up to 600°C. For the CS-PEO blended mat, it can be seen that the thermal stability has been increased so that ≈80% weight loss takes place at 600°C. The thermal stability is decreased by ≈15% in the CS-PEO/PLA/CS-PEO mat with a three-layered structure when compared with that of CS-PEO. Also, addition of 4% GT extract leads in increasing the weight loss by ≈15% compared with that of the CS-PEO/PLA/CSPEO mat. Considering the weight loss values, it can be concluded that strong intermolecular hydrogen bonding and electrostatic interactions between the functional groups of the two CS and PEO polymers with GT extract can produce a strong network leading to lower ash formation (greater weight loss) in the mat containing 4% GT with respect to that of the mat without GT extract. It is noteworthy that usually happens in the TGA analysis that addition of organic materials results in enhancement of weight loss, but addition of inorganic compounds causes decrease in the weight loss of the sample [48,50,52].

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The DSC profiles of the PLA polymer and CS-PEO, CS-PEO/PLA/CS-PEO, and CS-PEO-4% GT/PLA/CS-PEO-4% GT nanofibrous mats in Fig. 22.4C indicate that there are two peaks around 65°C and 125°C for the CS-PEO while at about 65°C and 150°C for the PLA, CS-PEO/PLA/CS-PEO, and CS-PEO-4% GT/PLA/CS-PEO-4% GT. The first peak can be attributed to the melt temperature (Tm) of PLA polymer, but the second peak may be correlated to the evaporation of water and acetic acid entrapped within the polymeric chains [60–62]. Additionally, there exists an endothermic peak in all DSC diagrams between 290°C and 360°C as well as another peak in the diagram of CS-PEO at 410.34°C that may be related to the degradation of polymer chains and GT herbal extract components. The TGA/DSC results exhibit that all of the three-layered CS-PEO/PLA/CS-PEO mats have similar degradation behaviors and they are thermally stable up to ~450°C.

Water uptake In order to obtain insight into the swelling behaviors of the fabricated CS-PEO-GT/ PLA/CS-PEO-GT mats containing 0%–4% of GT extract, their water uptakes were carried out at 25°C in three media including acidic (pH = 4), neutral phosphate buffer saline (PBS, pH = 7.4), and alkaline (pH = 13) solutions. The plots of swelling percents after 3 days are displayed in Fig. 22.5A. Since all of the nanofibrous mats possess a three-layered architecture and the swelling occurs at the mat surface, the CS-PEO layers are in contact with water molecules. The reason for the swellings of the mats can be attributed to the H-bond interactions of the amino NH2, acetamido CH3C(O) NH and hydroxyl OH groups of the CS, the CO, OH groups of PEO, and OH, CO, 0% GT 1% GT 2% GT 3% GT 4% GT

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Fig. 22.5  (A) The swelling diagrams of CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT extract. The (B) digital photographs of the swelled mats after 3 days in pH = 4.

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CO moieties of GT extract constituents with the aqueous medium. It is apparent from Fig.  22.5A that in all of the three acidic, neutral, and alkaline environments, the swelling is increased by adding 1% and 2% of GT to the CS-PEO layers but it is decreased by further GT addition (3% and 4%). The highest enhanced swelling due to the presence of 2% GT may be attributed to the increased H-bond interactions of the CS, PEO, and GT functional groups with the aqueous medium. Also, the decreased swelling by incorporation of 3%, 4% GT in the CS-PEO mat may be related to the highly enhanced H-bond interactions of CS, PEO, and GT with each other leading to reducing their interactions with the water molecules. For all of the CS-PEO-GT/PLA/CS-PEO-GT mats containing 0%–4% of GT, the swelling is the highest in an acidic medium, but decreases in PBS; the least water uptake is measured in the alkaline environment. The reason for such a behavior lies in the protonation of the amino NH2 and amido CH3C(O)NH groups of CS to form ammonium cations which enhance swelling of the mats [40,42,43]. The water uptake in the PBS buffer takes place only because of the H-bond formation between water and functional groups of polymers/GT components. In the alkaline medium, the OH− anions do not intend to greatly interact with the polymers; consequently the swelling degree is decreased. It is noteworthy that all of the mats maintain their integrity in the three acidic, neutral, and alkaline environments except for the mat incorporated with 2% GT with the highest swelling percent which indicates gelation of the nanofibers after 3 days only in the acidic medium (it is also stable in pH = 7.4 and 13). The digital photographs and SEM micrographs of the mats (containing 0%–4% of GT) suspended in acidic medium (pH = 4) for 3 days are presented in Figs. 22.5B and 22.6 in which it is evident that the mat with 2% GT has the highest water uptake and its texture has almost been destroyed and changed to a gel. Also, these figures confirm that the swelling changes in the order of 2% > 3% > 1% > 4% > 0%.

Water contact angle The water contact angle analysis is usually achieved to recognize the hydrophilicity of a mat surface. The pictures provided on top of each column in Fig. 22.7A reveal water droplets placed on the surface of each mat indicating a droplet are more dispersed (smaller angle) on a more hydrophilic mat but more spherical (it not dispersed) on a less hydrophilic mat surface (bigger angle). It has been established that CS and PEO are both hydrophilic polymers which are soluble in acidic and aqueous environments. As a result, in order to apply the CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats for biomedical applications, appropriate mats must be produced which are compatible with the aqueous medium. As can be seen in Fig. 22.7A, the contact angle is decreased by addition of 1%, 2% of the GT extract to the CS-PEO layers but it is increased by further adding the GT (3% and 4%) indicating the lowest surface wettability for the mat containing 0% GT while the highest hydrophilicity for the mat loaded by 2% of the GT extract. The highest hydrophilicity for the mat including 2% GT in the CS-PEO mat may be related to the increased H-bond interactions between the functional groups of the CS, PEO, and GT with the water molecules but the increased lipophilicity by incorporation of 3%, 4% GT may be described by the highly enhanced H-bond interactions

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Polysaccharide Carriers for Drug Delivery

Fig. 22.6  The SEM images of the swelled CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT extract after 3 days in pH = 4.

of CS, PEO, and GT with each other leading to the decrease in their interactions with the water molecules. Comparing these results with those obtained from the water uptakes reflects that they are consistent with each other and the water contact angle data support the swelling experiments. According to these results, it can be concluded that the CS-PEO-GT/PLA/CS-PEO-GT mat loaded by 3% GT extract is preferred over other mats for the biomedical applications because of its improved hydrophilic property.

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80

Contact angle ()

60

40

20

0 0% GT

1% GT

2% GT

3% GT

4% GT

Sample

(A)

Cumulative release (%)

70

35 1% GT 2% GT 3% GT 4% GT

0 0

(B)

200

400

600

Time (h)

Fig. 22.7  (A) The water contact angles and (B) the drug (GT) release profiles for the of CSPEO-GT/PLA/CS-PEO-GT mats containing 0%–4% of GT extract during 25 days in the PBS buffer.

Release of GT extract from the nanofibrous mats In order to decide which nanofibrous mat is more appropriate for wound dressing applications, the release of GT extract from CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% GT was examined during 25 days in the PBS buffer and the results are plotted in Fig. 22.7B. It is observed that the release of the GT is sharply increased within the first ~20 h (burst release) but after that the

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GT is released gradually until about 25 days after which the release rate becomes almost constant. Furthermore, as expected, the amount of GT extract released at the same time is higher from the fiber with greater amount of GT than those including a lower amount of GT. This result validates that the CS-PEO mat containing 4% GT is preferred over the mats incorporated with 0%–3% GT for biomedical applications.

FT-IR and UV-visible spectra The FT-IR spectra of GT extract and CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT are presented in Fig. 22.8. The spectrum of GT extract reveals a broad band at 3419 cm−1 which is related to the OH groups of its constituents (catechins, Fig. 22.1A) and also to the OH groups of water which is possibly adsorbed by the GT powder. The two peaks at 2924 and 2853 cm−1 are assigned to the CH stretching vibrations. The bands at 1697 and 1634 cm−1 correspond to the vibrational frequencies of carbonyl groups and the bands between 1556 and 1238 cm−1 are attributed to the stretching vibrations of double bonds (such as CC) but those that appeared at 1143, 1092, 1034, 872, 822, 762 and 608 cm−1 are related to the vibrations of CO and CC singe bonds. The FT-IR spectrum of the mat without GT illustrates a broad band between 3100 and 3400 cm−1 which is due to the stretching vibrations of NH and OH groups present in the CS and PEO polymers. A band at 2874 cm−1 is related to the vibrations of CH groups. The two bands with moderate intensities at 1554 and 1405 cm−1 are attributed to the vibrations of amide CH3C(O)NH groups of CS. The intense bands at 1067 and 1026 cm−1 plus a weak band at 653 cm−1 are correlated to the vibrations of CO, CN and CC singe bonds. In the FT-IR spectra of CS-PEO-GT/PLA/CS-PEO-GT mats containing 0%–4% GT, similar to the spectra of GT and the mat without GT, broad bands exist between 3100 and 3400 cm−1 correspond to the stretching vibrations of amine and hydroxide groups. Also, the bands at about 2800–2900 cm−1 are assigned to the CH stretching vibrations whereas bands near 1750, 1550, and 1450 cm−1 are characteristics of amide I, II, and III, respectively. The bands at approximately 1360, 1380, and 1260 cm−1 are attributed to the CH3 symmetrical deformation modes, and those at about 1100 and 1000 cm−1 are indicative of the CO, CN and CC stretching vibrations. The three weak bands at 750, 650, and 550 cm−1 are due to the wagging of CS saccharide structure. It is obvious that the number of peaks and their intensities are reduced upon increasing the GT extract weight percent in the CS-PEO layers reflecting the formation of greater H-bonds between the NH, CO, CO, OH groups of CS, and PEO with the OH, CO, CO moieties of GT components leading to disappearing/decreasing the bands corresponding to these functional groups of the polymers. These results on the FT-IR spectra are consistent with that reported in the literature [40,41]. The UV-visible spectra of the GT extract and CS-PEO-GT/PLA/CS-PEO-GT mats containing 0%–4% GT are given in Fig. 22.9. It is observed that the extract has a broad peak in the range of 200–270 nm. The pure mat without GT extract shows a sharp peak around 220 nm plus a broad peak between 200 and 400 nm. Also, all of the mats

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Fig. 22.8  The FT-IR spectra of GT extract and CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT extract.

containing 1%–4% GT reveal sharp bands at about 230 nm as well as near 200–400 nm which can be attributed to the n → π* and π → π* transitions of the GT extract components (catechins) and the polymers. Furthermore, all of the spectra of the mats loaded by the GT are identical, but the intensities of their peaks are enhanced by increasing the GT amount within the mats.

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Polysaccharide Carriers for Drug Delivery

Absorbance

7

GT extract 0% GT 1% GT 2% GT 3% GT 4% GT

3.5

0 200

400

600

Wavelength (nm)

Fig. 22.9  The UV-visible spectra of GT extract and CS-PEO-GT/PLA/CS-PEO-GT nanofibrous mats containing 0%–4% of GT extract.

Antibacterial activity The antibacterial properties of CS-PEO-GT/PLA/CS-PEO-GT mats containing 0%, 3% and 4% of GT extract were evaluated against both Gram-positive Staphylococcus aureus and Gram-negative E. coli (Fig. 22.10). Two different dilutions of each bacterial sample were used for the antibacterial experiments including 1:5000 and 1:50,000 dilutions. The mat without GT herbal extract reveals some antibacterial activity which is due to the antimicrobial effect of CS. Indeed, the antibacterial activities of CS-based materials have frequently been proved [40,42,43,47,52]. The nonantimicrobial control has 1 × 108 colony forming unit (CFU)/mL, but it is decreased in the case of CS-PEO/PLA/CS-PEO sample to 1.24 × 107 and 1.95 × 107 CFU/mL for its antibacterial potencies against Staphylococcus aureus and E. coli, respectively. The mats containing 3%, 4% GT illustrate higher antibacterial effects compared with the mat which is free of GT extract and this phenomenon can be attributed to the release of GT extract into the medium to kill the bacteria. The counted numbers of viable E. coli and Staphylococcus aureus bacteria (in more diluted solutions) using the mats with 3%, 4% GT contents are 3.91 × 106, 1.12 × 106 CFU/mL and 2.13 × 106, 0 CFU/mL, respectively indicating nearly 96%, 99% and 98%, 100% antimicrobial efficiencies. These results confirm that the antibacterial effects are greater against Gram-positive Staphylococcus aureus microorganism. Indeed, it has been found that although the Gram-positive bacteria have a thicker peptidoglycan layer in their cell wall, they are more susceptible to chemicals than Gram-negative ones because they contain impermeable lipid-based bacterial outer membrane. The possible mechanism proposed for the antibacterial activity of the CS containing compounds is based on the damaging interaction of the polycation (protonated amino groups, NH3+) with the negatively charged surfaces of bacteria, resulting in loss of membrane permeability, cell leakage, and finally cell death [44,45]. Consequently, performing functions such as carrying out processes necessary for respiration are prevented and finally the bacterial cell will die. Finally, it can be concluded that both of the mats containing 3% and 4% GT are suitable for biomedical applications such as wound dressing for regeneration of

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Laboratory controls

CS-PEO-3% GT/PLA/CS-PEO-3% GT

CS-PEO/PLA/CS-PEO

CS-PEO-4% GT/PLA/CS-PEO-4% GT

(A)

Laboratory controls

CS-PEO-3% GT/PLA/CS-PEO-3% GT

CS-PEO/PLA/CS-PEO

CS-PEO-4% GT/PLA/CS-PEO-4% GT

(B) Fig. 22.10  The digital photographs showing the antibacterial activities of CS-PEO-GT/PLA/ CS-PEO-GT nanofibrous mats containing 0%, 3%, and 4% of GT extract against (A) Gramnegative Escherichia coli and (B) Gram-positive Staphylococcus aureus bacteria.

v­ arious nerves because they are biocompatible and have high antibacterial properties. It is noteworthy that in many surgical applications, especially dirty operations or those involving the implantation of devices, one of the most important criteria is protection of the body from infection.

22.3 Hydroxyethylcellulose drug delivery systems Hydroxyethylcellulose (HEC) is one of the most significant commercially accessible cellulose derivatives [63]. It has been used as a thickener, coating, and stabilizer

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Polysaccharide Carriers for Drug Delivery

in numerous applications. Much attention is also given to its possible usage in hydrogel synthesis. The preparation of both chemical and physical gels is extensively investigated on the cellulose derivatives but in the case of HEC the research focus is on chemical gelation. The free radical cross-linking which is initiated by irradiation, gamma-radiation, electron beam, and UV irradiation using photoinitiators are some of the methods that are employed to initiate the reaction [64]. Indeed, the development of innovative materials using renewable organic resources by means of new processes is of note and merits attention from both academic and industrial researchers. Therefore, many natural polymers, particularly polysaccharides, have been employed to fabricate superabsorbent hydrogel materials which are favorable candidates for application in the biomedical field due to their exceptional advantages including biodegradability, biocompatibility, nontoxicity, and having biological functions [65]. Basically, cellulose is the most plentiful polymer in nature and its analogs are commonly utilized as a result of their solubility in water, compatibility with tissues and blood, and low cost. Recently, HEC/acrylic acid (AAc) copolymer gels with superabsorbent characteristics were fabricated from aqueous solutions through radiation-initiated crosslinking and the influence of the AAc amount on gel properties was examined under diverse synthesis conditions [66]. The partial replacement of the cellulose derivative with AAc enhanced the gelation and led to greater gel fraction and minor water uptake even in very low concentrations (1%–5%). The AAc affected the electrolyte sensitivity of the hydrogels so that pure HEC gels were unaffected by the ionic strength of the solvent whereas the water uptake of HEC/AAc gels was reduced with the salt concentration. The sensitivity was dependent on the AAc ratio. Several pH sensitive cryogels were developed using two biodegradable natural polymers—CS and HEC—by cryogenic treatment of semi-dilute aqueous solutions and UV-triggered cross-linking in the frozen state [67]. The N,N′-methylenebisacrylamide enabled the creation of a polymer co-network and improved both the mechanical strength of cryogels and gel fraction yield. All HEC-CS cryogels were opalescent sponge-like materials which rapidly released/uptook water because of their open porous structure. The HEC-CS cryogels were tested as drug delivery systems for a highly water soluble drug (metronidazole). The release profile of metronidazole showed initial burst release (about 40% of the drug was released during 30 min) and subsequently sustained metronidazole release happened for the next 8 h. Some pH-sensitive HEC/ hyaluronic acid (HA) complex hydrogels containing isoliquiritigenin were produced for application as transdermal delivery systems to treat skin lesions created due to pH imbalance [68]. HA has promising skin compatibility and HEC can act as a scaffold to form hydrogels with diverse HCE:HA mass ratios. It was found that the HECHA13 (HEC:HA mass ratio was 1:3) had optimum rheological and adhesive characteristics thus it was used to probe the drug release efficacy as a function of pH (the efficiency was more than 70% at pH = 7). The antimicrobial effects against Propionibacterium acnes were performed in order to exploit the pH-sensitive nature of HECHA13. It was shown that at pH 7, HECHA13 containing isoliquiritigenin could inhibit the growth of Propionibacterium acnes. Moreover, HECHA13 exhibited outstanding permeability of isoliquiritigenin into the skin, which entered mainly through the hair follicles.

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22.4 Carboxymethyl cellulose drug delivery systems Carboxymethyl cellulose (CMC) is a cellulose derivative possessing numerous carboxymethyl groups on its cellulose backbone and is extensively employed as a natural ingredient to prepare hydrogels for several applications such as wound dressing, drug delivery, tissue engineering, and plant breeding [69]. In particular, a new class of hydrogels denoted as superabsorbent hydrogels fabricated using natural polymers such as cellulose and its derivatives are established for biomedical applications [70]. Several ecofriendly hydrogel membranes were prepared using carboxymethyl cellulose (CMC) for application in wound healing and skin repair [71]. The CMC hydrogels were achieved by two functionalization degrees (0.77 and 1.22) and chemically cross-linked with citric acid (CA) and they were blended with polyethylene glycol (PEG, 10 wt%). Moreover, PEG had a significant role in the creation of a hybrid polymeric matrix. The CMC-based hydrogels were cytocompatible according to the in vitro cell viability responses of >95% for human embryonic kidney cells (HEK293T) which were utilized as a model cell line [71]. It is well known that lack of antibacterial efficacy, poor water oxygen, and vapor permeability, as well as unsatisfactory mechanical properties are drawbacks of existing wound dressing materials [72]. On the other hand, hydrogels absorb wound exudates because of their robust swelling abilities and provide a cool feeling and a wet environment. Thus in order to solve these deficiencies, flexible nanocomposite hydrogel films were developed by incorporation of zinc oxide impregnated mesoporous silica (ZnO-MCM-41) as a drug nano-carrier into the CMC hydrogel using citric acid as cross-linker in order to avoid the cytotoxicity induced by commonly used cross-linkers [72]. The drug delivery and antibacterial activity of the hydrogel films were examined using tetracycline (TC) antibiotic and a sustained TC release was measured that would effectively decrease the need to exchange the bandage. Cytocompatibility of the hydrogel films was evaluated in adipose tissue-derived stem cells (ADSCs) that indicated the cytocompatibility of CMC/ZnOMCM-41. In another study, it was aimed to enhance the skin permeability, the loading efficiency of transdermal drugs, and their sustained release properties by preparing biodegradable nanocomposite hydrogels using in situ synthesized AuNPs on CMC cross-linked with poly(methacrylic acid) [73]. The in vitro release of diclofenac sodium and diltiazem hydrochloride proposed that in situ incorporation of AuNPs on cross-linked CMC-PEG improved the diffusion ability of hydrogel and released the drugs in a controlled manner (~79% for diclofenac sodium and ~85% for diltiazem hydrochloride were released after 3 days).

22.5 Gelatin drug delivery systems Some naturally derived polymers which are utilized in biomedical applications can be classified as proteins and polysaccharides [74]. Several frequently used protein-based polymers are gelatin, elastin, collagen, and albumin. The fundamental benefits of the developing polymeric materials using natural polymers for biomedical purposes are

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Polysaccharide Carriers for Drug Delivery

related to the strong cell-polymeric material interactions and biodegradability of polymers [75]. Recently, the gelatin-tamarind gum (TG)-based filled hydrogels were prepared for drug delivery applications and, for this purpose, three diverse types of carbon nanotubes (CNTs) were added to the dispersed TG phase of the hydrogels and the drug release experiments, both iontophoretic and passive, proved that the non-Fickian diffusion of the drug was predominant throughout its release from the hydrogel matrices. Moreover, the hydrogels were cytocompatible with human keratinocytes [76]. It is recognized that gelatin is a nonimmunogenic and biodegradable biopolymer but gelatin materials have restricted mechanical features (particularly in aqueous solutions) and high hydrophilicity leading to challenging delivery of a high load and sustained release of hydrophobic drugs [77]. In order to solve these problems, gelatinbased hydrogel films were added to a low concentration of PLGA and aminolyzed PLGA (PLGA-NH2) to deliver hydrophobic drugs. Indomethacin was cross-linked with diverse concentrations of 1,6-hexamethylene diisocyanate in DMSO solution. The biocompatible gelatin/PLGA-NH2 hydrogel films showed high indomethacin encapsulation efficiencies (approximately 84.8%) with low burst release behaviors, sustained indomethacin release up to 96 h, and the highest accumulative release (76.6 ± 3.7%) among the hydrogel films. In another work, several carboxymethyl chitosan (CMCS)/gelatin hydrogels were obtained by radiation cross-linking and a preclinical test was carried out by implantation model and full-thickness cutaneous wound model in Sprague-Dawley rats in order to preliminarily assess the biodegradability, biocompatibility, and influences on healing [78]. In order to estimate the function on wound healing, the hydrogels were used on rather large full-thickness cutaneous wounds (Φ3.0 cm). In comparison with the control groups, the hydrogel group exhibited considerably higher percentage of wound closure on days 9, 12, and 15 postoperatively, which was in agreement with the meaningfully thicker granulation tissue on days 3 and 6. Despite progress in production of materials containing therapeutic agents for diabetic wound treatment, impaired wound healing by unsatisfactory chemotactic responses has still remained an important issue [79]. Accordingly, in a recent study, horseradish peroxidase (HRP) enzyme-catalyzed hydrogels were developed as sprayable bioactive wound dressing materials that could deliver cell-attracting chemokines for diabetic wound healing [79]. It was assumed that topical administration of chemokines by the hydrogels could recover wound healing through prompting recruitment of the endogenous cells. The chemotactic cytokines (interleukin-8 (IL-8) and macrophage inflammatory protein-3a (MIP-3a)) were readily loaded into the hydrogels during the in situ gelling process and wound healing effectiveness of chemokine-loaded hydrogels was examined in streptozotocin-induced diabetic mouse model. Also, the bioactivities of MIP-3a and IL-8 released from hydrogel matrices were preserved. In vivo transplantation of chemokine-incorporated hydrogels assisted cell infiltration to the wound area and supported wound healing with improved re-epithelialization/neovascularization and enhanced collagen deposition relative to no treatment or the hydrogel alone. Thus the hydrogels considerably supported wound healing with quicker wound closure, neovascularization, and thicker granulation. Some gelatin/CMCS hydrogels were produced by radiation-induced-cross-linking at ambient temperature and the

Polysaccharide hydrogel films/membranes

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physicochemical and biological characteristics of the hydrogels were explored to assess their abilities as wound healing materials [80]. The enzyme degradation rate of the hydrogels could be controlled by regulating the procedure which was matched to the healing rate of the wound. Besides, the gelatin/CMCS hydrogels stimulated cell attachment and rapid growth of fibroblasts on their surfaces. In another attempt, lupeol entrapped chitosan-gelatin hydrogel (LCGH) films were fabricated by a solution casting process through blending CS and G solutions using glycerol as a plasticizer and then cross-linking with glutaraldehyde [81]. The antioxidant assay justified that lupeol and LCGH film had exceptional antioxidant properties by scavenging both radicals at fixed increasing rate which was increased with time because of stable release of lupeol. It was found that the antibacterial activity of lupeol in LCGH film was preserved. The cell viability test using MTT assay on NIH/3T3 fibroblast cells exhibited that the CGH film obviously had acceptable cell viability and nontoxicity. The in vivo wound healing potential of gels and films fabricated using cuttlefish skin gelatin (CSG) including aqueous henna extract (50 and 500 μg/mL) was recently investigated [82]. It was indicated that adding aqueous henna extract to the gelatin gels and films improved the antioxidant effect in a dose-dependent manner. In the wound healing experiment, topical application of gelatin hydrogels/films supplemented by aqueous henna extract, on the wound area in a rat model, considerably increased wound healing efficacy and assisted to avoid inflammation damage, compared to the control and CICAFLORA-treated groups. Moreover, all formulations enhanced the antioxidant status of treated rats which was proved by improved antioxidant enzymatic activities. The increased activities of superoxide dismutase, catalase, and gluthatione peroxydase in wound tissues displayed that bioactive compounds in aqueous henna extract, gel, and film could promote wound healing by reducing the damage produced by reactive oxygen species. In another study, biocompatible hybrid hydrogel scaffolds were produced using methacrylate-functionalized high molecular mass CS with gelatin-A which was photocross-linked with UV radiation. It was found that the hydrogels were biocompatible because the live cell viability responses of human embryonic kidney (HEK293T) cells were >95% [83]. A bilayer membrane was fabricated as a wound dressing material using a commercial polyurethane wound dressing as the outer layer and electrospun gelatin/keratin nanofibrous mat as the inner layer [84]. The cell attachment tests exhibited that adhered cells were spread better and were migrated deeper to the gelatin/keratin mat compared with that to the gelatin mat. The animal experiments revealed that in comparison with the bilayer membrane without keratin, gauze, and commercial Comfeel wound dressing, the bilayer membrane provided a much greater number of blood vessels and greater reduction in the wound area after 4  days, and superior wound repair after 14 days with a thicker epidermis and higher number of freshly formed hair follicles. In order to fabricate wound dressings, gelatin composite hydrogel fibers were developed by gel-spinning using PEG6000 as the modifier and dialdehyde carboxymethyl cellulose (DCMC) (a perfect cross-linking agent) was utilized to fix the hydrogel fibers [85]. The biological characteristics of the hydrogel fibers were evaluated to be used as wound dressings. Incorporation of DCMC effectively enhanced the mechanical properties, blood compatibility, and enzymatic stability

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Polysaccharide Carriers for Drug Delivery

of the hydrogel fibers. Cross-linking by DCMC decreased the swelling degree of the hydrogel fibers which was favorable to prevent losing their mechanical properties.

22.6 Gellan gum drug delivery systems Gellan is an exopolysaccharide produced by Pseudomonas elodeac which has negative charges when it is dissolved in water [86]. It is used as the gelling agent due to being biodegradable, biocompatible, with low toxicity, and creates transparent hydrogels having net negative charges. Fabrication of the polyelectrolyte complex between gellan and CS has been investigated by researchers. Recently, drug-polymer electrostatic interactions have been used to develop a unique ternary CS-ibuprofen-gellan nanogel to control transdermal delivery of ibuprofen [87]. The ternary nanogels were achieved by combining electrostatic nanoassembly and ionic gelation methods. The ex vivo delivery of ibuprofen onto and across the skin was assessed based on particular drug release parameters. CS improved the skin penetration, permeability, and the rate of transdermal ibuprofen release by a factor of 4, which was governed by the extent of ibuprofen-CS ionic interaction and its concentration. Some controlled release systems for ciprofloxacin were achieved by synthesis of gellan gum derivatives containing quaternary ammonium groups through grafting N-(3-chloro-2-hydroxypropyl)-trimethyl ammonium chloride to the primary hydroxyl groups of gellan by nucleophilic substitution reaction [88]. The in vitro transdermal release of ciprofloxacin from loaded particles were performed on rat skin in isotonic PBS solution (pH = 7.43) and it was found that ciprofloxacin was released up to 24 h which confirmed quaternized gellan-CS particles were promising in controlled release systems for topical dermal purposes [88]. It is known that improving skin permeation is an important factor to develop novel transdermal drug delivery formulations and this is principally significant for nonsteroidal anti-inflammatory drugs. In order to achieve this goal, semisolid gel and solid hydrogel film formulations comprising gellan gum as a gelling agent were produced and the influences of diffusion enhancers (isopropyl alcohol, dimethyl sulfoxide, and propylene glycol) was examined on carriage of the nonsteroidal anti-inflammatory diclofenac sodium [89]. After 24 h, greater amount of diclofenac was released from the solid hydrogel film compared with that of new and commercial diclofenac formulations. Entrapment of temperature-responsive nanogels in the solid hydrogel film yielded temperature-activated extended diclofenac release. When diclofenac was entrapped into the temperature-responsive nanogels added to the solid hydrogel film, its transport was minimal at 22°C, however, it was enhanced sixfold when the temperature was increased to skin surface temperature of 32°C. Apigenin (APN) was extracted from M. alba leaves and screened by in  vivo wound models (diabetic and dead space) in rats and the Apigenin containing hydrogels were fabricated using gellan gum-CS (GGCS) with PEG as a cross-linker and characterized by AFM, entrapment efficiency, swelling, and drug release tests [90]. Other characteristics of hydrogels were assessed by wound healing activity experienced through wound contraction, dried granuloma weights, collagen content, and

Polysaccharide hydrogel films/membranes

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a­ ntioxidant p­ erformance. The APN-GGCS hydrogels exhibited higher wound healing influence in both diabetic and normal wound tissues with noteworthy antioxidant activity.

22.7 Guar gum drug delivery systems Guar gum is the powder endosperm obtained from the Cyamopsis tetragonolobus seeds (a leguminous crop). The endosperm is composed of a complex polysaccharide named galactomannan which is a polymer made up of d-mannose and d-galactose. The polymer has numerous hydroxyl groups that enable it to form hydrogen bonds in water leading to forming a highly viscose and condense solution. Due to its condensing, binding, emulsifying and gelling characteristics, extensive pH stability, rapid solubility in cold water, film forming capability, and biodegradability, it finds applications in diverse industries such as agriculture, food, paper, textile, explosives, hydraulic fracturing, bioremediation, cosmetics, drug delivery, pharmaceuticals, and medicine. Guar gum backbone is comprised of linear (1 → 4)-β-d-mannopyranosyl chain units along with branch points of α-d-galactopyranosyl units which are linked by (1 → 6) bonds. The ratio of mannose to galactose units has been found in the range of 1.6:1 to 1.8:1 but, in another study, the ratio was estimated to be 2:1 [91]. Since the uronic acid is absent in the structure of guar gum, it is dissimilar to many plant gums and mucilages. It was mentioned that since the transdermal administration of diltiazem hydrochloride was not very effective due to slow penetration of drug through the skin, gold nanoparticle (GNP) was used to improve the permeation to the skin [60]. A transdermal device was prepared using a polyelectrolyte complex that was reinforced with nanogold-nanocellulose (GNP-NC) composites. The films containing hybrid filler were made with diverse loading contents and the loading effect was assessed on tensile properties, drug encapsulation efficiency, thermal stability, water vapor permeability, as well as skin permeability. Furthermore, the influences of storage time and temperature on the drug release behavior were examined. Also, the fabricated devices were tested for in vivo skin adhesion and irritation in human subjects, environmental fitness, and cell viability in order to evaluate their possible application in pharmaceutical areas.

22.8 Acacia gum delivery systems Acacia gum (AG) or gum Arabic is a dried exudation which is achieved from the branches and stems of Acacia Senegal or other types of acacia (fam. Leguminosae). Acacia gum primarily contains higher molecular weight polysaccharides as well as their magnesium, calcium, and potassium salts which produce galactose, arabinose, rhamnose, and glucuronic acid upon hydrolysis [61]. AG is water soluble which forms solutions with various concentrations that are not very viscous. It is a nondigestible

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food ingredient that is widely used in the food and pharmaceutical industries. Also, it has robust anti-oxidant, nonhemolytic, hemostatic, and antibacterial nature [62]. Thus by keeping in mind the significance of using polysaccharide gums in wound repair, it was attempted to discover the antioxidant nature of AG by development of hydrogel wound dressing with improved wound healing potential [92]. The polymers were fabricated using AG/polyvinyl pyrollidone/carbopol and essential biomaterial properties of wound dressings including wound fluid absorption, bioactive assessment, hemo-compatibility, water/gass/microbial permeability, mechanical characteristics, bio-adhesion, drug release, and histology of wound healing were investigated. The hydrogel wound dressings had antioxidant, nonhaemolytic, and mucoadhesive nature. Another work was intended to achieve an ideal wound dressing biomaterial possessing anti-inflammatory and antibacterial properties without any toxicity to the host cells but offering the most healing activity [93]. Thus zinc oxide nanoparticles (ZnONPs), having antimicrobial effect and being able to improve wound healing, were used to prepare ZnONPs-loaded‑sodium alginate-acacia gum (SAAG-ZnONPs) hydrogels through cross-linking hydroxyl groups existing in the polymeric chains of sodium alginate and gum acacia with the aldehyde group of gluteraldehyde. Then, the wound healing influence of sodium alginate/gum acacia/ZnONPs was examined. The SAAGZnONPs hydrogels displayed a healing influence with a low concentration of ZnONPs employed on sheep fibroblast cells. Hence, the findings proposed that high concentrations of ZnONPs were toxic to cells, whereas SAAG-ZnONPs hydrogels considerably decreased the toxicity and maintained the favorable antibacterial and healing effects.

22.9 Pectin drug delivery systems Pectin (AP) is a polysaccharide composed of galacturonic acid units and known as a structural constituent of plant cell walls [94]. It is found in roots, leaves, seeds, flowers, and fruits. Pectin-based hydrogels have been used in pharmacy because of their biodegradability, biocompatibility, pH-responsive properties, and nontoxicity [95]. Also, the AP drops the blood cholesterol level, decreases the glucose uptake, and has anti-tumor activity. Naturally, the galacturonic acid units on the AP-based hydrogels are ionized leading to their repulsion and therefore more efficient swelling and releasing of drugs through diffusion. Consequently, the AP-systems can satisfactorily deliver drugs to the colon [96]. Asiatic acid (AA)-pectin and chloroquine (CHQ)-pectin patches were prepared in order to estimate the effect of AA transdermal delivery on physicochemical variations, parasitaemia and associated pathophysiology of Plasmodium berghei-infected Sprague Dawley rats [97]. To reach this goal, a topical one-off AA (5, 10, and 20 mg/kg)- or chloroquine (CHQ)-pectin patch was used on the shaven dorsal neck region of Plasmodium berghei-infected Sprague Dawley rats (90–120 g) on day 7 subsequent to infection. Drinking and eating habits, weight alterations, malaria effects, and parasitaemia were compared among animal groups for 21 days. Also, all animals developed stable parasitaemia (15%–20%) by day 7 and AA doses could considerably inhibit parasitaemia so that the most efficient patch was AA 5 mg/kg. The AA and CHQ exhibited bimodal

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time-spaced peaks but CHQ patch needed an extended time for the clearance of parasitaemia. Moreover, AA affected the parasitaemia suppression and bio-physicochemical properties in a dose dependent way. Considering the administered drug dosage, AA revealed much greater effectiveness compared with CHQ confirming that AA might be a beneficial antimalarial drug. Magnetic pH- and thermo-responsive microgels were prepared using pectin maleate, N-isopropyl acrylamide, and Fe3O4 NPs [98]. Curcumin was incorporated to the microgels and release tests were conducted. A slow sustained curcumin release was realized under external magnetic field condition. The loaded curcumin presented bioavailability, stability, and superior solubility compared to the free curcumin. Moreover, the cytotoxicity experiment displayed that magnetic microgels without curcumin suppressed the growth of Caco-2 cells. In another study, pristine and blended pectin hydrogels and poly(3methoxydiphenylamine) conductive polymer/hydrogel were prepared containing ferrous chloride (FeCl2) and citric acid as the cross-linking agents and ibuprofen drug through the solution casting method [99]. The in vitro ibuprofen release from the hydrogels was examined in buffer solution (pH = 5.5) at 37°C for 48 h. In the case of both cross-linkers, the drug diffusion coefficient was improved by decreasing the cross-linking ratio. By applying an electric field of 5 V, the drug diffusion coefficients of the FeCl2 and citric acid hydrogels were boosted by a factor greater than two because the electro-repulsive interactions occurred between ibuprofen and the negatively charged electrode. Furthermore, the poly(3-methoxydiphenylamine) added to the pectin hydrogels as the drug encapsulating host increased the diffusion coefficient and decreased the total release time. Polymeric hydrogels were created through cross-linking of poly(methylvinylether-co-maleicacid) (PMVE/MA) and pectin (PE) for application in pharmaceutics [100]. Several blend hydrogel patches were fabricated as transdermal delivery formulations and the influences of diverse preparation factors were examined in relation to their rheological and pharmaceutical characteristics [101]. The pectin and gelatin mixtures were used to produce patches having variable properties. Furthermore, diverse modalities for drug loading were studied regarding testosterone drug homogeneous distribution. The abilities of the transdermal patches were examined in terms of reliability and reproducibility by obtaining the in vitro drug release profiles. Since natural polymeric hydrogel films are suitable materials for wound dressing application, some simvastatin loaded cross-linked alginate-pectin hydrogel films were produced through ionic cross-linking which exhibited a sustained slow release of simvastatin and the in vitro cytotoxicity test proved the nontoxicity of the hydrogel films [102]. Pectin-CS hydrogel drug carriers were prepared in order to evaluate the mechanism governing the release pattern and the drug release assay established sustained release of three drugs including curcumin, progesterone, and mesalamine during 24 h under physiological conditions [103]. Several dual responsive hydrogels were achieved using pectin, poly((2-dimethylamino)ethyl methacrylate), and bis[2-methacryloyloxy] ethyl phosphate cross-linker through free radical polymerization [104]. The hydrogels were used to encapsulate 5-fluorouracil anticancer drug and to produce silver NPs. The 5-fluorouracil incorporated hydrogels were utilized for the in vitro drug release in both simulated gastric and intestinal environments (pH = 1.2 and pH = 7.4) at two

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temperatures (25°C and 37°C). The antibacterial effects of nanocomposite hydrogels were evaluated against four bacteria: E. coli, Klebsiella pneumoniae, B. cereus, and S. aereus and it was indicated that the maximum activity occurred against Grampositive S. aereus bacterium followed by B. cereus, moderate activity was detected for Gram-negative E. coli, and the minimum was obtained using K. pneumoniae.

22.10 Hyaluronic acid drug delivery systems Hyaluronic acid (HA) is a naturally occurring linear polysaccharide made up of repeating N-acetyl-d-glucosamine and d-glucuronic acid units in which the monosaccharides are bound together by successive β-1,3 and β-1,4 glycosidic bonds [105]. The pKa value for the HA carboxyl groups is 3–4, thus the functional groups are mainly ionized at pH 7.4 and it is a polyanion under the physiological conditions named hyaluronan. HA is produced with varied molecular weights that range from 20 to 4000 kDa and its different molecular weights are also applied in pharmaceutical formulations [106]. The HA chains have random-coil conformations in solution and they are very hydrophilic being surrounded with water molecules linked by hydrogen bonds. This conformational feature along with its high molecular weights means the HA solutions are viscous and elastic. In nature, HA primarily exists in the extracellular matrix of connective tissues and it is the most plentiful in the vitreous body of the eye. Furthermore, HA has vital roles in intracellular functions due to the fact that it can affect cell proliferation and migration; moreover it can modulate intracellular signaling [107]. For pharmaceutical applications, HA is known as a nonimmunogenic, nontoxic, biocompatible, and biodegradable polymer. HA can be chemically functionalized through conjugation and cross-linking reactions using its functional groups including carboxylic acid and hydroxyl groups [108]. Since HA has promising physico-chemical and biological properties, it has been used extensively in numerous biomedical applications such as treatment of ocular and plastic surgery, osteoarthritis, tissue engineering, and drug delivery. Moreover, it has been used as a carrier for receptor-mediated drug targeting in cancer therapy and delivery of peptide, protein, and nucleotide therapeutics along with delivery of imaging compounds because of its ability to distinguish receptors whose expression is enhanced in different diseased cells. Recently, hydrogel nanoparticles (37 nm size) were synthesized by cross-linking between HA and PEG and used as efficient carriers for transdermal drug delivery [109]. The hydrogel NPs exhibited diverse capabilities for penetration into the skin layer depending on the type of dispersion medium. The in vitro skin penetration tests confirmed that the hydrogels dispersed in an oil or an oil composition more effectively penetrated into the skin of albino guinea pigs compared with those distributed in an oil-in-water emulsion. The preclinical tests established that the hydrogel NPs dispersed in the oil or the oil composition were fairly entered to the skin layer of albino guinea pigs. Linear or cross-linked HA hydrogels as semisolid drug delivery systems were compared for deep HA penetration into the skin; their hydration, irritation, in vitro, and in vivo skin penetration were examined and it was established that the smaller

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HA particles formed by cross-linking diffused more to a synthetic membrane and entered further into the human epidermis and living animal skin compared with linear HA which did not penetrate [110]. The properties of chemically cross-linked hydrogels made using HA-HEC were examined for efficient transdermal delivery of isoliquiritigenin [111]. The in vitro skin penetration experiment revealed that the system considerably enhanced the isoliquiritigenin delivery to the skin. It was intended to increase the analgesic influences of rutaecarpine and evodiamine by a microemulsion hydrogel for the transdermal codelivery and to evaluate HA hydrogel for microemulsion entrapment [112]. Hence, a microemulsion was prepared using ethyl oleate as the oil core to enhance the solubility of the alkaloids. The effect of increasing permeation of rutaecarpine and evodiamine in ME-Gel by microemulsion led to attaining 2.60- and 2.59-fold higher transdermal fluxes relative to the hydrogel control. The HA hydrogel containing microemulsion displayed favorable skin biocompatibility but efficient ME-Gel co-delivery of rutaecarpine and evodiamine towards the skin improved the analgesic influence in mouse pain models in comparison with the hydrogel.

22.11 Chondroitin sulfate drug delivery systems Chondroitin sulfate (ChS) is a biopolymer which is ample in the extracellular matrix (ECM) of tissues in the body and has essential biological and structural functions. ChSs are a particular group of glycosaminoglycans (GAGs) which contain repeating sulfated disaccharides comprising N-acetylgalactosamine. Also, the ChS chains are categorized to types A, C, D, E, K, and H which are called ChSs A (C-4-S), C (C-6-S), D (C-2,6-S), E (C-4,6-S), K (C-3,4-S), and H (IdoAa1–3Gal-NAc(4S,6S)), with the most common ChSs being C-4-S and C-6-S. CSs have shown vital effects in wound healing [113]. Diverse ChSs can stimulate diverse and occasionally even contradictory cell responses. For instance, ChS-A prevents axonal growth but ChS-E (CS-E) reveals a reverse result [114]. Moreover, ChSs are capable of partially regulating other biological functions such as coagulation, inflammation, stem cell niche creation, enzymatic activity, apoptosis, and complement activity. Thus addition of a suitable ChS to an adhesive for a special application improves tissue regeneration by the fabricated biomaterial [115]. Attempts have been made to describe the interactions occurring between cells and different kinds of ChSs [116]. Recently, a ChS-PEG adhesive hydrogel was developed for biomedical applications [117]. The adhesive strength of the hydrogel to the cartilage tissue was 10 times greater compared with that of fibrin glue. The ChS-PEG hydrogel resulted in minimum inflammatory response along with the in vitro enzymatic degradation once it was implanted subcutaneously in a rat model. In another work, gamma-ray irradiation was used to achieve biocompatible hydrogels for application in skin tissue engineering [118]. The hydrogels were fabricated using natural polymers including hyaluronic acid (HA) and ChS as well as the synthetic polymer poly(vinyl alcohol) (PVA). Also, HA/ChS/PVA hydrogels with different compositions

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were obtained by gamma-ray irradiation method to be used as artificial skin alternatives. The human keratinocyte cell cultures contacted with the HA/ChS/PVA extracts demonstrated >92% cell viability. The human keratinocyte growth on HA/ChS/PVA hydrogels was slowly improved depending on the culture time. It was illustrated that after 7 days, the human keratinocyte cells in the case of all HA/CA/PVA hydrogels presented >80% viability relative to the human keratinocyte culture control group. It was attempted to fabricate an elastic and tough hydrogel scaffold using ChS and HA by their conjugation with tyramine (TA) [119]. Hydrogels were achieved through mixing ChS-TA and HA-TA in the presence of horseradish peroxidase (HRP) and H2O2 which exhibited fatigue resistance against cyclic compression. Confocal laser scanning microscopy proved that mesenchymal stem cells that were encapsulated into the hydrogels had appropriate cell viability. Some HA/ChS/poly(acrylic acid) (PAAc) hydrogels were fabricated via gamma-ray irradiation without using other initiators or cross-linkers in order to prepare a biocompatible hydrogel to be used in skin tissue engineering [120]. The hydrogels including ChS, HA, glycosaminoglycans, and PAAc as a synthetic ionic polymer were produced as artificial skin alternatives by gamma-ray irradiation using simultaneous cross-linking and free radical copolymerization. The in vitro drug release rates of the hydrogels were considerably affected by the swelling plus the interactions occurred between the ionic groups existing in the hydrogels and the ionic drug molecules. The cytotoxicity tests using human keratinocyte cells cultured on the HA/ChS/PAAc extracts illustrated that all of the hydrogels had comparatively extraordinary cell viabilities of >82% and did not lead to any noteworthy adverse signs on cell viability.

22.12 Collagen drug delivery systems Collagen, HA and elastin are the main structural constituents of the dermis. Moreover, collagen is the central component of connective tissues with types I and III being the most plentiful collagen produced by skin fibroblasts [121]. Collagen is an exceptional triple-helical structural protein and the major part in the extracellular matrices occurring in all multicellular animals. Collagen molecules include three α-chains entangled in the well-known collagen triple-helix which is stabilized through intra- and inter-chain hydrogen bonds formed between nearly constantly repeating Gly-X-Y-sequence (X is typically proline and Y is commonly hydroxiproline) [122]. Furthermore, high molecular weight collagen shows structural integrity to the skin which forms a dense network in the dermis and affords structural support to the epidermis. Low molecular weight collagen can enhance the diameter and density of collagen fibers, increases HA creation, and stimulates protecting against UV-A radiation [122]. Collagen is extensively applied in biomedical areas such as wound dressing, hemostatic, tissue engineering scaffolds, and drug delivery systems. Collagen peptide is hydrolyzed from collagen which exhibits well antioxidant activity and is mainly valuable for wound healing purposes. Applications of collagen in colloidal drug delivery systems consist of sponges for wounds/burns, scaffolds, mini-pellets and tablets for protein delivery,

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nanoparticles for gene delivery, gel formulation containing liposomes for sustained drug delivery, and collagen matrices [123]. A semi-conductive nanocomposite as an electrically controlled drug delivery system was prepared and for this purpose hydrolyzed collagen (a natural plentiful polypeptide) was modified using polycaprolactone [124]. In situ aniline polymerization inserted conductive nanofibers in the hydrogel matrix. Moreover, in vitro conductivitystimulus hydrocortisone drug release was examined. The MTT assay confirmed that the conductive and nonconductive hydrogels had no cytotoxicity. It was mentioned that the improvements in skin related health subjects are the subject of much attention due to the effectiveness of microemulsion in transdermal and dermal delivery of active pharmaceutics/ingredients [125]. Hence, the water-in-oil microemulsion capability was evaluated to combine the two natural polymers including collagen and HA having low and high molecular weights. It was found that depending on the collagen molecular weight and surface activity, the droplet size of the microemulsion was changed. In another work, microbial transglutaminase (MTGase) was utilized as a catalyst in order to perform grafting of the collagen peptide (COP) molecules onto the amino groups of CS to synthesize water-soluble CS-collagen peptide (CS-COP) compounds [126]. The dynamic blood clotting experiment confirmed that the hydrogels had suitable blood coagulation ability. The hydrogel biocompatibility was assessed using NIH-3T3 cells with the MTT assay [126]. Since nanofibrous scaffolds containing traditionally common medicine as a wound dressing material can stop infection and promote wound healing, collagen (COL) extracted from marine fish skin was used to coat the poly(3-hydroxybutyric acid) (P)-gelatin (G) nanofibrous scaffold by a bioactive Coccinia grandis extract (CPE) developed with electrospinning [127]. The collagen coated scaffold (PG-CPE-COL) was applied onto the experimental wounds of rats to analyze its wound healing efficacy. The collagen and CPE in the nanofibrous scaffolds supported the wound healing and thus decreased the inflammation triggered by the cyclooxygenase-2 and inducible nitric oxide synthases. The nanofibrous scaffold affected the expression of different growth factors including vascular epidermal, endothelial, and transforming growth factors. Furthermore, the PG-CPE-COL scaffold improved the collagen synthesis and accelerated reepithelialization.

22.13 Alginate drug delivery systems Alginate is a natural anionic polysaccharide which can form hydrogels [128]. It is extensively used because of its biodegradability, biocompatibility, low toxicity, low cost, and gelling characteristics. Commercially existing alginates are usually achieved using brown algae (Phaeophyceae) such as Laminaria digitata, L. hyperborea, L. japonica, Macrocystis pyrifera and Ascophyllum nodosum. Also, alginates are produced by bacterial biosynthesis from Pseudomonas and Azotobacter having more recognized chemical structures and physical properties in comparison to the seaweed derived alginates [129]. The chemical structure of alginate comprises (1,4)-linked β-dmannuronate and α-l-guluronate residues which can form homopolymeric or

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alternating sequences in the polymeric chains. The contents of M and G blocks, their distributions in the chains, and the blocks’ lengths highly govern the alginate chemical and physical properties and its gelling ability. Alginate is broadly applied to prepare gels in the food industry due to it having exceptional gel formation features in aqueous solutions in the presence of several multivalent cations. The cross-linking and gelation of alginate occurs through the exchange of sodium ions by multivalent cations. For instance, alginate hydrogels can be obtained with external gelation using calcium ions that are used as cross-linking agents. The cations interact and bind with guluronate blocks in the alginate chains to create the gel network [128]. This cross-linked hydrogel is beneficial for application in controlled release of pharmaceutical/bioactive compounds. Furthermore, alginate gel is commonly used to treat numerous wounds as a result of its extraordinary elasticity, water content, permeability, and capability to offer a humid environment in the wound area [130]. Particularly, calcium alginate hydrogels are employed in wound healing purposes because of both the hemostatic characteristic of the calcium ions and the capacity of the gel to aggregate erythrocytes and platelets. Moreover, the porosity of the alginate hydrogels leads to the immobilization/entrapment of therapeutic compounds (such as growth factors and drugs) which will be released to the wound bed. This method is used in wound infection treatments that are commonly required in the initial phases of the course of wound healing. Recently, smooth, flexible and thin interpenetrating network hydrogels were developed using sodium alginate (SA) and poly(vinyl alcohol) (PVA) through a solvent casting process for transdermal delivery of prazosin hydrochloride which is an anti-hypertensive drug [131]. The in vitro drug release test was accomplished from excised rat abdominal skin and it was found that the interpenetrating network hydrogel membranes extended the drug release to 24 h but SA and PVA membranes rapidly released the drug. Moreover, the skin histopathology and skin irritation assays confirmed that the interpenetrating network hydrogel membranes caused lower irritation and could be used for skin care applications [131]. It is known that alginate and Aloe vera are natural materials that are extensively examined and utilized in biomedical fields. Thus thin hydrogel films containing alginate and Aloe vera gel with diverse ratios (95:5, 85:15, and 75:25 v/v%) were fabricated through cross-linking by 5% calcium chloride and it was found that the films had satisfactory mechanical features for skin application whereas the solubility test established their insolubility after 24 h immersion in distilled water [132]. It was attempted to obtain a useful controlled release transdermal delivery system for selegiline by means of thermosensitive hydrogels [133]. The copolymers of alginate (A) and Pluronic F127 (PF127) were employed to prepare thermogels through both physical blending (A + P) and chemical grafting (AP). The skin penetration across nude mouse skin and porcine skin confirmed that the thermogels had sustained release of selegiline with AP showed the most sustained permeation. AP hydrogels exhibited linear permeation properties for the transdermal delivery of selegiline. Because photocrosslinkable biomaterials are favorable in tissue engineering because they are capable of being injected and creating in situ hydrogels in a minimally invasive way, photocrosslinked alginate hydrogels (HP-A) were developed [134]. An affinity-based growth factor delivery system was achieved via incorporation of hepa-

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rin to the HP-A which permitted prolonged and controlled release of therapeutic proteins. The release of growth factors from these affinity-based platforms was sustained for 3 weeks without any early burst release with the released growth factors preserving their biological activities. Implanting bone morphogenetic protein-2 containing HP-A hydrogels promoted moderate bone formation nearby the implant periphery. Notably, morphogenetic protein-2 loaded HP-A hydrogels caused considerably more osteogenesis compared with morphogenetic protein-2 loaded unmodified hydrogels so that 8 weeks after implantation, 1.3- and 1.9-fold superior calcium content and peripheral bone creation were measured for the BMP-2-incorporated HP-A hydrogels relative to the BMP-2 containing unmodified alginate hydrogels. In another study, a nanohybrid interpenetrating network hydrogel was fabricated using laponite:polyvinyl alcohol-alginate (LAP:PVA-A) having controllable biological, physical, and mechanical features for application in wound healing [135]. The hydrogels exhibited excellent biocompatibility against fibroblast and MG63 cells. The MTT assay confirmed that fibroblast proliferation was considerably greater on 0.5 wt% LAP:PVA-A in comparison with the PVA-A. Additionally, the hydrogels supported hemostasis which would be useful in the wound healing process. A covalently crosslinked hydrogel was achieved using N,O-carboxymethyl chitosan (N,O-CMCS) and oxidized alginate (OA) in order to be used as a drug delivery system [136]. The in vitro and in vivo biocompatibility and cytocompatibility of the hydrogel were assessed and it was found that the hydrogel had suitable cytocompatibility with NH3T3 cells after 3 days incubation. Also, the hydrogel did not produce any cutaneous reaction during 72 h of subcutaneous injection subsequent to slow degradation and adsorption by the time. The extraction of hydrogel had almost 0% of hemolysis ratio, demonstrating its appropriate hemocompatibility. A biodegradable antibacterial OA-CMCS composite hydrogel was fabricated which was filled with microspheres in order to be employed in drug delivery and wound healing applications [137]. Some tetracycline hydrochloride incorporated gelatin microspheres were produced and added to the OA-CMCS hydrogel to achieve the desired dressing. Among other formulations, the hydrogel containing 30 mg/mL of microspheres displayed more appropriate stability and mechanical property for wound healing. It was exhibited that the loaded tetracycline hydrochloride was released in a sustained manner from the hydrogel contrary to pure hydrogel and microspheres. The powerful inhibition of Staphylococcus aureus and E. coli bacterial growth confirmed that the hydrogel dressing, particularly the hydrogel loaded with 30 mg/mL gelatin microspheres comprising tetracycline hydrochloride, could find an auspicious future in bacterial infection treatments. The effect of halloysite nanotubes as an inorganic support was estimated on the biological activity and release rate of the vancomycin antibiotic loaded alginate-based dressings [138]. The halloysite nanotubes were incorporated with ~10 wt% of vancomycin antibiotic and subsequently they were encapsulated into the alginate and gelatin/alginate hydrogels. It was found that the material functionalized with aliphatic amine considerably prolonged the vancomycin release from the hydrogels compared with that obtained using silica. For comparison, after 24 h, the antibiotic amount released from silica was 70% but the vancomycin drug release was 44% from halloysite in alginate discs. Addition of gelatin caused even

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further extended sustained drug release. Only the immobilized drug loaded in alginate revealed excellent antimicrobial potency against different bacteria because the inhibition zones were larger than those of the standard discs against the Enterococci and Staphylococci bacteria. The in vivo toxicity test was carried out using Acutodesmus acuminatus and Daphnia magna and it was exhibited that the hydrogels did not influenced the living organisms.

22.14 Smart hydrogels as drug delivery systems Stimuli-responsive hydrogels are widely investigated and identified as smart materials that are sensitive to pH, temperature, light, and magnetic field [139]. Such materials are excellent candidates for application in drug delivery because their release behaviors can be modified exactly through an external stimulus [140]. Hence, several stimulus–responsive membranes have so far been developed by means of various polymers including biopolymers such as chitosan and alginate. Using thermo-sensitive macromolecules, compounds having a multiple and independent sensitivity to ambient variations were prepared. As an example of temperature sensitive polymers, poly(N-isopropylacrylamide) (PNIPAAm) is one of the most extensively examined thermo-sensitive polymers displaying extraordinary hydration-dehydration alterations in aqueous medium in response to rather small temperature changes around its lower critical solution temperature (LCST), about 30°C [141]. Also, it is very significant that PNIPAAm has very low toxicity via elimination by glomerular filtration. In fact, transdermal drug delivery systems (TDDS) are prepared using thermo-responsive polymers (TRPs) to control the drug amount delivered through skin and to make an on-demand drug delivery. TRPs are smart polymers demonstrating volume phase transition and consecutively an abrupt solvation state variation at a definite temperature named LCST which is related to the TRPs that become insoluble on heating or UCST that corresponds to the TRPs which are soluble on heating. It is noteworthy that the compatibility of the polymeric hydrogel with skin is essential to avoid skin irritation and it is dependent to different factors including skin hydration, pH of skin, and trans-epidermal water loss. It is known that the pH of the normal skin surface is measured in the range of 5.4–5.9 which regulates as the skin barrier and prevents skin infections. Therefore, it is vital for the transdermal carrier hydrogel to preserve pH between 5 and 7 to cause no irritation to the skin. At pH values of 8 and higher, it harms the skin surface through enhancement of the trans-epidermal water loss and leads to skin irritation [142]. Recently, a smart alginate membrane hydrophobically modified by sodium palmitate and biomineralized with CaHPO4 was prepared containing indomethacin drug and it was found that the drug release from the membrane was both pH- and thermo-responsive [143]. The drug release was about 60% during 12 h but the alginate membrane revealed that the release was >90%. It was shown that the biomineralized hydrophobic alginate membrane delayed the penetration of the encapsulated drug into the solution and decreased the drug release rate. Smart auto-degradable microspheres were developed using calcium alginate/high methoxylated pectin comprising an alginate lyase (A) from Sphingobacterium multivorum and levofloxacin [144]. The microspheres were achieved through ionotropic

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gelation method using inactive form of A at pH 4.0. It was found that incubation of microspheres in PBS and Tris-HCl buffers at pH 7.4 exhibited the influence of ionchelating phosphate on the matrix erodability which proposed an inherently A activation via changing the pH near the neutral value. Encapsulating the levofloxacin drug to the A containing microspheres presented 70% efficiency with 35% improvement in the antimicrobial activity against Pseudomonas aeruginosa biofilm. The levofloxacin release from microspheres was not different at acidic pH however it was changed at neutral pH when using A. The mammalian CHO-K1 cell cultures proved that the matrix was cytotoxic and the levofloxacin toxicity was decreased after its encapsulation. It was aimed to design a heat triggered transdermal drug delivery system (TDDS) using a thermo-responsive polymer including poly(N-vinyl caprolactam) (PNVCL)based gel [145]. Using this TDDS, whenever pain is felt by patients, they will be able to administer a pulse of drug by simple application of heat pad on the TDDS. The triggered drug delivery was examined via loading both hydrophobic etoricoxib and hydrophilic acetamidophenol drugs. To evaluate the drug release in response to pH and temperature, in vitro drug release tests were accomplished at three diverse temperatures (25°C, 32°C, and 39°C) and two pH values (5.5 and 7.0). The drug release profiles exhibited superior release for both the drugs at 39°C (above LCST) and pH 5.5. The in vitro skin permeation of both the drugs was carried out in rat abdominal skin which displayed greater drug release when the skin temperature was higher (39°C). Additionally, skin permeation by the hydrophobic drug was enhanced compared with that of the hydrophilic drug. The in vivo biocompatibility assay for the CP hydrogel on rat skin ascertained the gel biocompatibility. Since the inventive drug delivery systems based on smart hydrogels for localized on-demand drug delivery are very significant, a smart UV-cross-link cross-linkable and thermo-responsive CS hydrogel was prepared for NIR-triggered localized ondemand drug release [146]. The CS hydrogel was synthesized through its grafting onto poly(N-isopropylacrylamide), acetylation of methacryloyl groups and incorporation into photothermal carbon. The methacryloyl groups acted as the UVcross-link cross-linkable unit which led to CS gelation upon UV irradiation. The poly(Nisopropylacrylamide) served as the thermo-responsive unit that caused the CS hydrogel to show temperature-triggered volume shrinkage and reversible swelling/de-swelling behaviors. The CS hybrid hydrogel incorporated into the photothermal carbon displayed individual near-infrared triggered volume shrinkage (~42%) in response to temperature rise which was created by near-infrared laser radiation. Also, the doxorubicin drug release rate was enhanced to almost 40 times greater than that of nonirradiated hydrogels. Smart composite hydrogels (SCHs) involving CS microspheres physically incorporated into a thermo-responsive hydrogel were achieved in order to assess their ability to load and long-term drug release [147]. The CS microspheres were used because they are pH-sensitive and could electrostatically attach the oppositely charged salicylic acid drug molecules. The salicylic acid showed a sustained release from SCHs which was consistent with temperature and pH. Salecan is known as a water soluble polysaccharide created by a salt-tolerant strain Agrobacterium sp. ZX09 and poly(dimethylaminoethyl methacrylate) (PDMAEMA) is a pH, ionic strength and thermal multi-sensitive polymer having antibacterial p­ roperty.

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Thus a semi-interpenetrating polymeric network (semi-IPN) hydrogel was fabricated using salecan and PDMAEMA which was simultaneously sensitive to pH, ionic strength, and temperature [148]. The cytotoxicity assessment proved that the hydrogel was noncytotoxic to COS-7 cells. The in  vitro release rate of insulin protein as the model drug was improved by enhancement of the hydrogel swelling ratio. Furthermore, when the temperature was greater than the LCST of PDMAEMA, the hydrogel was shrunk to discharge additional drug molecules. Higher release rate and release amount were measured in acid medium (pH 1.2) compared with at pH 7.4. It is recognized that transdermal protein delivery is a beneficial way of protein therapy and dermal vaccination but the method is limited due to the low permeability of the stratum corneum [149]. Consequently, a transdermal delivery system was developed to increase protein permeability to the skin. For this purpose, a transparent gel patch was produced that was obtained using polysaccharides with gold nanorods on the gel surface and fluorescein isothiocyanate-modified ovalbumin inside. The gel patch was located on mouse skin in order to be in contact with the gold nanorods and irradiated with a continuous-wave laser to heat the gold nanorods and raise the skin temperature to 43°C which improved the translocation of the isothiocyanate-modified ovalbumin to the skin. Because penetration increasing materials can be coated on biodegradable polymeric nanogels containing cytotoxic drugs for topical application, they are capable of augmenting the chemotherapeutic efficacy of skin cancers [150]. Therefore, in a study, the in  vitro and ex  vivo chemotherapeutic ability of double-walled biodegradable PLGA-CS nanogel incorporated with 5-fluororuacil drug and coated with eucalyptus oil was examined by its topical application on the skin. Hence, 5-fluororuacil drug was entrapped into the PLGA core through the solvent evaporation method and then coated with cationic CS for ionic interactions with anionic skin cancer cells. The surface coating by 1% eucalyptus oil was done to increase the penetration efficiency of the nanogel to the stratum corneum. The ex vivo screening of porcine skin was used to evaluate the permeation ability of the nanogels. A sustained drug release was measured over 24 h. A low hemolysis (3.31%) along with short prothrombin time and activated partial thromboplastin time of 13.5 and 33 s, respectively, proved the astonishing biocompatibility of the nanogels. The cytotoxicity test by MTT assay using human keratinocyte cell line verified noticeable cytocompatibility of the nanogels. The ex vivo assessment on porcine skin exhibited effective and constant flux of 5-fluororuacil from the nanogels to the skin but the histology of the porcine skin demonstrated greater penetration efficacy of eucalyptus oil coated PLGA-CS double-walled nanogels.

22.15 Conclusion Polysaccharides are natural biocompatible polymers containing modifiable functional groups which makes them ideal compounds for biomedical applications. Also, they are nontoxic, biodegradable, and biocompatible having film forming ability. Among various polysaccharides, chitosan, hydroxyethylcellulose, carboxymethyl cellulose, gelatin, gellan gum, guar gum, acacia gum, pectin, hyaluronic acid, chondroitin sulfate, collagen, alginate, and smart hydrogels have been used extensively in ­transdermal

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drug delivery. For instance, thiocolchicoside incorporated chitosan/hyaluronan transdermal films containing different weight ratios of hyaluronan to CS showed that when less hydration of the polymeric matrix occurred, the drug diffusion and its penetration to the skin were lower, confirming such formulations could be used as exceptional platforms for transdermal drug delivery. In our research work, biocompatible antimicrobial nanofibrous mats were designed with a three-layered structure and produced by an electrospinning technique in which the chitosan-polyethylene oxide formed the first and third layers and contained 0%–4% of green tea ethanolic extract while the polylactide composed the middle second layer. It was established that the mat loaded by 3% GT was the best tissue/device (among the mats fabricated) for biomedical applications such as wound dressing due to it being biocompatible, having appropriate antibacterial properties, hydrophilicity, stability and tensile strength/elongation at break. Recently, antimicrobial pH-sensitive hydroxyethylcellulose/hyaluronic acid hydrogels incorporated with isoliquiritigenin were developed as transdermal delivery systems for the treatment of skin lesions generated due to pH imbalance. In another study, some antibacterial ZnONPs-loaded‑sodium alginate-acacia gum hydrogels exhibited a healing effect using a low concentration of ZnONPs employed on sheep fibroblast cells confirming the hydrogels considerably diminished the toxicity and preserved the auspicious antibacterial and healing efficacy. As an example of pectin hydrogels, magnetic pH- and thermo-responsive microgels containing curcumin drug were achieved using pectin maleate, N-isopropyl acrylamide, and Fe3O4 NPs. Currently, hydrogel nanoparticles are produced through cross-linking of HA and PEG which are employed as effective carriers in transdermal drug delivery. In a recent effort, a chondroitin s­ ulfate–PEG adhesive hydrogel was prepared and it was found that it caused minimum inflammatory response and in vitro enzymatic degradation when it was implanted subcutaneously in rats. Recently, a smart alginate membrane was hydrophobically modified by sodium palmitate, biomineralized with CaHPO4, and loaded by indomethacin drug; it was indicated that the drug release from the membrane was both pH- and thermo-responsive. The hydrophobic biomineralized alginate membrane postponed the drug penetration to the solution and declined the drug release rate. Thus it can be concluded that polysaccharide hydrogel films/membranes are valuable candidates for transdermal delivery of therapeutics.

Acknowledgments The financial support of this work by the Research Office of Amirkabir University of Technology (Tehran Polytechnic) is gratefully acknowledged.

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