**2. Mechanism of drug release form hydrogels**

The release of drugs from hydrogels can be achieved by different mechanisms such as swelling/deswelling, diffusion, and chemical mechanism. As previously mentioned, hydrogels are three-dimensional crosslinked polymeric networks that swelled in the presence of water. The crosslink can be physical (hydrophobic interactions, electrostatic interactions and hydrogen bonding) or chemical (covalent bonding) and is responsible of the network structure of the hydrogel. Such networks display open spaces, the size of which is referred to as the mesh size of the hydrogel [9]. Importantly, the mesh size of the hydrogels is one of the main parameters that affect how drugs diffuse through the hydrogel network, being dependent on polymer and crosslinker concentrations, as well as external stimuli. The gelation of hydrogels

*Sustained Drug Release from Biopolymer-Based Hydrogels and Hydrogel Coatings DOI: http://dx.doi.org/10.5772/intechopen.103946*

by polymerization means leads to network irregularities and polymer polydispersity upon formation, and as a result, the mesh size is usually heterogeneous. A number of approaches exist to determine the mesh size [9].

When the mesh is larger than the drug (rmesh/rdrug > 1), the drug release process is dominated by diffusion. Small drug molecules migrate freely through the network, and diffusion is largely independent of the mesh size. The diffusivity, D, in this situation depends on the radius of the drug molecule (rdrug) and the viscosity of the solution (η) via the Stokes–Einstein equation Eq. (1) [10]:

$$D = \frac{RT}{6\pi\eta r\_{dry}}\tag{1}$$

where R is the gas constant and T is the absolute temperature.

When the mesh size is close to the drug size (rmesh/rdrug ≈ 1), the effect of steric hindrance on drug diffusion becomes relevant. Finally, for an extremely small mesh size and/or very large drug molecules (rmesh/rdrug < 1), strong steric hindrance immobilizes the drugs and it remains physically entrapped inside the network, unless the network degrades or the mesh size expands in response for example, to external stimuli.

Several methods accompanied by mathematical model development have been created in parallel to hydrogel technology, in order to predict drug release from the network. The drug release fitting models (i.e. the zero order equation; the first order equation; the Higuchi's equation; the Korsmeyer-Peppas' equation; the Hixon-Crowell's equation, the Weibull equation, among others) are the most abundant, however, they are not predictive but simple mathematical fitting equations. In the last years, mechanistic and statistical models are growing quite fast. Mechanistic models combining the mass transport with the system mechanics developed with a "fully coupled" approach considers the influence of the mass transport on the mechanics as well as the opposite, which makes this approach the only candidate to produce reliable first-principle models.

Statistical models, are receiving a lot of attention due to the consensus of the regulatory authority and the possibility to predict the hydrogels behavior, in the analyzed design space, regardless the complicate phenomenology, with quick and inexpensive experimental designs [11]. Recently, Wu and Brazel developed a method for the simulation of water uptake profile and drug release from homogeneous hydrogels. This model successfully predicted the initial burst release observed experimentally [12]. Sheth et al. developed a mathematical and computational model using time snapshots of diffusivity and hydrogel geometry data measured experimentally as inputs to predict release profiles of two model proteins of varying molecular weights from degradable hydrogels [13].

### **3. Drug release from physical hydrogels**

Physical hydrogels are those formed by reversible and dynamic crosslinks grounded on noncovalent interactions. In this regard the network of physical hydrogels is reversibly held together by molecular entanglements, resulting from a dynamic competition between pro-assembly forces (for example, hydrophobic interactions, attractive electrostatic forces and hydrogen bonding) and anti-assembly forces (for example, solvation and electrostatic repulsion [3]. These interactions that occur in

this type of hydrogels are usually weak. However, they are numerous and contribute to the presence of complex behaviors.

Polyampholytes may also be used to construct physical hydrogels, with randomly dispersed cationic and anionic groups. The randomness leads to a wide distribution of strengths: The strong bonds serve as structural crosslinks, imparting elasticity, whereas the weak bonds reversibly break and re-form, dissipating energy. Consequently, physical hydrogels have reversible liquid to solid transition, also called sol–gel transition, in response to different changes in environmental conditions such as temperature, ionic strength, pH, or others [14]. Since the interactions depend significantly on external stimuli, they allow hydrogels to be highly versatile concerning the environment, unlike covalently bonded materials [15].

Physical hydrogels can be engineered to undergo spontaneous biodegradation under physiological conditions, which constitutes another way of controlling the release of active molecules [16]. Degradation is typically mediated by hydrolysis [17, 18] or enzyme activity [19]. The erosion or loss of polymer mass through degradation, can take place simultaneously in the bulk or on the surface of the hydrogel. For a variety of hydrogels, the bulk and surface erosion can be tuned to obtain desirable release kinetics ranging from weeks to months. Bulk erosion occurs because of the permeability to water or degrading enzymes when the rate of diffusion of these agents is rapid compared to the rate of bond degradation. Surface erosion, in contrast, results when the rate of bond breakage is more rapid than the rate of enzyme or water diffusion from the exterior into the bulk of the gel [13].

Representatives of reversible physical hydrogels are the shear-thinning hydrogels which flow like low-viscosity fluids under shear stress during injection, but quickly recover their initial stiffness after removal of shear stress in the body [3]. Alginate hydrogels are shear-thinning, formed via electrostatic interactions between alginate and multivalent cations (for example, calcium and zinc). They can be readily injected via a needle after gelation in a syringe and have been used to achieve sustained local delivery of bioactive vascular endothelial growth factor (VEGF) in ischemic murine hindlimbs for 15 days [20, 21].

#### **3.1 Peptide based physical hydrogels**

Another example of physical hydrogels forming materials are the self-assembling peptide systems, where amino acid-based chains undergo the sol–gel transition without the need of any chemical crosslinking agent. This property makes them useful materials to safely in situ encapsulate living cells or sensitive drugs, among others. In addition, this peptide-driven self-assembly into physical hydrogels is highly specific, sourced mainly by the biorecognition of peptide segments scattered among the macromolecular chains. They form dynamic well-defined, hierarchically organized 3D structures with reversibility of the assembly and disassembly processes [22]. Another example are elastin-like polypeptides cross-linked via electrostatic interactions between their cationic lysine residues and anionic organophosphorus cross-linkers [23]. Non-covalent interactions between heparin and heparin-binding peptides and proteins can also be used to form hydrogels for growth factor delivery [24, 25].

Peptide self-assembly can also be achieved by taking advantage of interactions between metal cations and amino acid residues of the peptides. This was demonstrated with gelation of a β-sheet-rich fibrillar hydrogel with zinc ions [26].

*Sustained Drug Release from Biopolymer-Based Hydrogels and Hydrogel Coatings DOI: http://dx.doi.org/10.5772/intechopen.103946*

#### **3.2 Chitosan based physical hydrogels**

Additional interesting example of physical hydrogels for drug release applications are pectin-chitosan hydrogels, which showed to be thermo-reversible and capable of prolonging the release of three different model hydrophobic drugs: mesalamin, curcumin and progesterone. In vitro drug-release studies revealed that lower percentage of pectin in the hydrogel led to slower release rates owing to smaller mesh size arising from stronger interactions between the polyelectrolytes. Also, the release was slower when the total polymer concentration was higher. Finally, a slower release in PBS solution compared to HCl solution was attributed to the fact that at pH 7.4, both polymers are charged, with strong electrostatic forces and consequently, smaller mesh size. At the molecular scale, the polymer chains can possess abundant binding sites for the drugs. DSC and FTIR analysis exhibited some interactions between the drugs and both chitosan and pectin that can contribute to the prolonged release of the drugs [27]. Another in situ-gelling hydrogel was formed with a polyelectrolyte complex, which showed a sustained release of insulin and avidin proteins [28].
