Temperature-sensitive or thermo-responsive hydrogels and polymers have attracted great attention in different fields because some diseases manifest a temperature modification [28, 29]. In particular, they are characterized by a critical solution temperature (CST) around which the hydrophobic and hydrophilic interactions between the polymeric chains and the aqueous media brusquely change within a small temperature range bringing to the disruption of intra- and inter-molecular interactions and resulting in chain collapse or expansion. Typically, these polymer solutions possess a temperature above which one polymer phase exists, namely upper critical solution temperature (UCST), and below which a phase separation occurs with the formation of an hydrogel [30]. On the other hand, polymer solutions that appear as monophasic below a specific temperature and biphasic above it generally possess a so-called lower critical solution temperature (LCST). The LCST is a fascinating phenomenon found for various water soluble polymers which tend to phase-separate from solution upon heating. The most investigated temperature-responsive polymer featuring a LCST in water is the poly(N-isopropylacrylamide) (PNIPAAm). The LCST of PNIPAAm was of ~32 °C, proximate to the human body temperature. By altering the temperature of PNIPAAm solution in water, its solubility behavior can be reversibly changed from hydrophilic-soluble to hydrophobic-insoluble and thus can be induced externally (Fig. 12.2). Then, above the LCST, the polymer become increasingly insoluble leading to gel formation.

The phenomenon of transition from a solution to a gel is commonly referred to as sol-gel transition. The sol–gel transition of thermosensitive hydrogels can be experimentally verified by a number of techniques such as spectroscopy [31–33], differential scanning calorimetry (DSC) [31, 32] and rheology [33].

Other N-substituted polyacrylamides [34, 35], and further classes of polymers such as poly(oligoethyleneoxide-(meth)acrylate)s [36], poly(2-oxazoline)s [37], poly(N-vinylalkylamides), e.g. poly(N-vinylcaprolactam) (PNVC) [38], copolymers such as poly(L-lactic acid)-poly(ethylene glycol)-poly(L-lactic acid) (PLLA-PEG-PLLA) triblock copolymers [39], poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) (PEO–PPO–PEO) copolymers [40] and many natural polymers [41], e.g. polysaccharides and proteins, utilize temperature change as the trigger that determines their gelling behavior without any additional external factor. Researchers have used them alone or in combination to fabricate thermally responsive hydrogels with desired properties. Hydrogels based on these polymers have been interesting for biomedical uses as they can swell in situ under physiological conditions and provide the advantage of convenient administration.

Photo-sensitive polymers have received increasingly attention in recent years because light can be applied straightaway and under specific conditions with high accuracy. The light is both negligibly absorbed by both cells and tissue and greatly consequently by the polymers. This makes light-responsive polymers highly advantageous for applications in wide fields including drug release, biomaterials, and artificial tissues [54–60]. Photo-responsive polymers contain light-sensitive chromophore moieties. Without external stimuli, these polymers can maintain their structures but, upon light absorption, these moieties can be broken from the polymer chain that can be degraded into smaller molecular fragments. Generally, light-sensitive polymers, UV or visible light sensitive, are classified into two main categories that are the side-chain-type polymer containing the chromophores as lateral groups of the chain backbone and the main-chain-type structure possessing single or multiple photosensitive species covalently- or non-covalently connected to the backbone. Chromophores groups are azobenzene [11, 61], spiropyran [62, 63] and nitrobenzyl [64, 65] and a variety of azobenzene or spiropyran-containing photo-responsive polymers such as poly(acrylamide) (PAA) [41, 66] and PNIPAAm [43, 44].

Increasing research and development efforts have been dedicated to the field of electro-responsive polymers (ERPs) due to their advantages of precise control via the magnitude of the current, the duration of an electrical pulse or the interval between pulses [76, 77]. ERPs are promising candidate materials for the design of drug delivery technologies, especially in conditions where an “on‐off” drug release mechanism is required. Typical ERPs are naturally occurring polymers such as chitosan, alginate and hyalouronic acid are commonly employed to prepare electro-responsive materials. Major synthetic polymers that have been used include polythiophene (PT) or sulphonated-polystyrene (PSS), polyaniline, polypyrrole, ethylene vinyl acetate, and polyethylene.

In some cases, combinations of natural and synthetic polymers have been used. Most polymers that exhibit electro-sensitive behavior are polyelectrolytes and they undergo deformation under an electric field due to anisotropic swelling or de-swelling as the charged ions move towards the cathode or anode (Fig. 12.5). Greatest stress is felt by the region surrounding the anode and smaller stress near the vicinity of the cathode. This stress gradient contributes to the anisotropic gel deformation under an electric field [78, 79].

Neutral polymers that exhibit electro-sensitive behaviour require the presence of a polarizable component with the ability to respond to the electric field. A rapid bending of gel in silicon oil was observed in the case of lightly cross-linked poly(dimethylsiloxane)-containing electrosensitive colloidal SiO₂ particles.

These polymers can show swelling, shrinking or bending in response to an external field [80, 81] and they may be blended into responsive hydrogels in conjunction with the desired drug to obtain a patient‐controlled drug release system useful for neurotransmitters and vaccine delivery. The “on‐off” drug release mechanism can be achieved through the environmental‐responsive nature of the interpenetrating hydrogel‐EAP complex via (i) charged ions initiated diffusion of drug molecules due to an increase in osmotic pressure in the polymer; (ii) conformational changes that occur during redox switching of EAPs; or (iii) electro‐erosion. For example, when an electrochemical stimulus is applied to multilayer polyacrylamide films, the combined effects of H ions migrating to the region of the cathode and the electrostatic attraction between the anode surface and the negatively charged acrylic acid groups lead to shrinking of the film on the anode side [78, 82].

The hydrogel‐EAP composites include other implications such as, application towards biosensors, DNA hybridizations, micro‐surgical tools and miniature bioreactors and may be utilized to their full potential in the form of injectable devices as nanorobots or nano‐biosensors.

An innovative approach of exploiting ultrasound in drug delivery consists on the application of the ultrasonic waves directly at the polymers or the hydrogel matrix [87]. For the ultrasonically responsive polymers the release mechanism is driven by the cavitation. In particular, the ultrasound (ULS) energy generates both high and low pressure waves, resulting in an alternative growth and shrinkage of gas-filled microbubbles. These high-low pressure waves regulate the intermittent opening of the pores of the polymer, thus inducing the delivery of the respective drugs.

The most widely employed polymers for the ULS-responsive uses could be biodegradable or non-biodegradable. The biodegradables ones include polyglycolides, polylactides, bis(p-carboxyphen-oxy)alkane anhydride with sebacic acid. Instead, the non-biodegradables materials are ethylene-vinyl acetate copolymers or the poly(lactide-co-glycolide) microspheres, PHEMA hydrogels, the PEO-b-PPO-b-PEO micelles and the poly(HEMA-co-DMAEMA) hydrogels. The releasing agents are p-nitroaniline, p-amino-hippurate, bovine serum albumin and insulin. When exposed to ultrasound, these bioerodible polymers respond rapidly and reversibly. It is believed that the ultrasound also causes an increase in temperature in the delivery system, which allows the diffusion [88, 89]. ULS-responsive polymers are advantageous because they are non-invasive and capable of penetrating deep into the interior of the body and thus drug delivery can be focused and carefully controlled through a number of parameters including frequency, power density, duty cycles and time of application [90, 91]. Biologically the ultrasound action is related to the generation of thermal energy, perturbation of cell membranes, and enhanced permeability of blood capillaries [92].

The use of an oscillating magnetic field to modulate the rates of drug delivery from a polymeric matrix was one kind of method to achieve externally controlled drug delivery systems [107]. When planning a magnetic-responsive delivery system characteristics such as the magnetic properties of the carrier particles, field strength and geometry, drug/gene binding capacity and physiological parameters like the depth to target, the rate of blood flow, vascular supply and body weight need to be considered [108]. Magnetic targeting is based on the attraction of magnetic polymer carrier to an external magnetic field source that effectively traps it in the field at the target site and pulls it toward the magnet [109–113]. In order to maintain the magnetic loaded carrier at a specific location, the externally applied field must have a relatively strong gradient. When the drug is released from the magnetic matrix this is no longer responsive to the applied field.

Several carriers for magnetically enhanced drug delivery have been studied including nanoparticles with a magnetic core and a polymer shell and liposomes which have a magnetic core and an artificial liposomal shell are reported [114–119]. These materials may also be embedded in hydrogels which can carry drugs that are released upon heating [110, 120]. To produce highly efficient magnetic-responsive materials, the “doping” of polymer materials with magnetic nanoparticles (MNPs), made of inorganic matter (most often superparamagnetic iron oxide (SPION) Fe₃O₄ or g-Fe₂O₃, or ‘‘soft’’ metallic iron, but also “hard” magnetic materials e.g. Co, Ni, FeN, FePt, FePd…), appeared to be the more appealing and efficient solution.

Many hydrophilic polymers were proposed to host magnetic nanoparticles, in particular thermosensitive gels like poly(N-isopropylacrylamide) (PNIPAAm). However, PNIPAAm and other polymers exhibiting a LCST are generally less polar than polyvinyl acetate (PVA) and can be inefficient for trapping the MNPs due to the absence of hydrogen bonds and to a mesh size of typically 10–20 nm, i.e. larger than the size of superparamagnetic iron oxide MNPs (5–10 nm) [121]. To overcome this problems, strategies reported in the Fig. 12.7 concerning on a statistical copolymer network with a chelating comonomer such as 2-acetoacetoxyethylmethacrylate (AEMA) [122], a semi-interpenetrated network with alginate chains wrapping the MNPs [123] or a composite network of PNIPAAm and poly(ethylene glycol) (PEG) using PEG400–dimethacrylate as crosslinker of NIPAAm [124] were proposed.

pH sensitive polymers are materials which respond to the changes in the pH of the surrounding medium by varying their dimensions. Such materials increase its size (swell) or collapse depending on the pH of their environment (Fig. 12.8). This behavior is due to the presence of certain functional groups in the polymer chain such as acidic or basic groups that can either accept or release a proton in response to changes in environmental pH.

The pH is an important environmental parameter for biomedical applications, because pH changes occur in several specific or pathological sites.

The changes along the gastrointestinal tract from acidic in the stomach (pH = 2) to basic in the intestine (pH = 5–8) has to be considered for oral delivery of any kind of drug. Certain cancers as well inflamed or wound tissue exhibit a pH different from 7.4 as it is in circulation. For example chronic 7.4 and 5.4 and cancer tissue is also reported to be acidic extracellularly [135–137].

Consequently, this parameter can be exploited for a direct response at a certain tissue or in a cellular compartment. As mentioned above, pH responsive polymers consist of polymers with a large number of ionisable groups known as polyelectrolytes. Polyelectrolytes are classified into two types: weak polyacids and weak polybases. Weak polyacids accept protons at low pH and release protons at neutral and high pH [138]. Poly(acrylicacid)(PAAc) and poly(methacrylicacid) (PMAA) are commonly used pH-responsive polyacids [139, 140]. As the environmental pH changes, the pendant acidic group undergoes ionization at specific pH called as pKa. This rapid change in the charge of the attached group causes alteration in the molecular structure of the polymeric chain. This transition to expanded state is mediated by the osmotic pressure exerted by various stimuli responsible for controlling drug release from smart polymeric delivery systems [26–143]. pH responsive polymers typically include also chitosan [144], albumin [145], gelatin [146], poly(acrylic acid) (PAAc)/chitosan IPN [147], poly(methacrylic acid-g-ethylene glycol) [P(MAA-g-EG)] [148, 149], poly(ethylene imine) (PEI) [150], poly(N,N-diakylamino ethylmethacrylates) (PDAAEMA), and poly(lysine) (PL) [151, 152].