## 4 Drug Delivery
### Temperature-Responsive Drug Delivery  
Higher temperatures at disease sites provide another venue to target and control the release of encapsulated drugs from nanoparticles. In addition, heat can be applied from external sources in order to trigger the nanoformulation response. Polymers used in temperature-responsive systems typically feature a lower critical solution temperature (LCST) near the physiological temperature, above which a polymer becomes immiscible. The most common temperature-responsive polymer used with drug delivery systems, and the only one used in miktoarm star-based drug delivery systems is poly(N-isopropylacrylamide) (PNIPAM), due to its availability, biological compatibility, and well-studied LCST in the range of 30-35 \({}^{\circ}\)C [136]. In a very early example, an AB\({}_{2}\) (A = PNIPAM, B = poly(undecylenic acid) (PUA)) miktoarm star was synthesized that had an LCST of 31 \({}^{\circ}\)C, very close to accepted values for PNIPAM [27]. Upon loading prednisone acetate, a common anti-inflammatory drug, it was reported that at temperatures below LCST, even after 150 h, more than 80% of the drug remained intact. However, significantly faster release rates were seen above the LCST of PNIPAM. In a later study, similar AB\({}_{3}\) miktoarm star polymers (A = PMMA, B = PNIPAM) were synthesized [31], and below the LCST of PNIPAM, these formed spherical micelles in a size range of 50 nm, and with a CMC of 10 mg/mL. Prednisone acetate was stabilized in the PMMA cores of the micelles, and there was a total cumulative release of 55% at room temperature, compared to a 90% release above LCST. In such cases, the conversion of PNIPAM to a hydrophobic polymer above its LCST results in its collapse from the micellar corona into the core. As a result of the morphological transformation, an increase in the cumulative release of prednisone takes place [27; 31]. While the previous examples deal with the corona/core switch of PNIPAM, one could also stabilize its position in micellar coronae via crosslinking. Crosslinking the micellar shells of assembled AB\({}_{2}\) (A = PNIPAM, B = poly(L-lysine) (PLL)) miktoarm star polymers would not lead to micellar collapse following the temperature increase above the LCST of PNIPAM, and the effect on drug delivery would be more subtle [38]. In such a system, shell crosslinking of drug-loaded micelles was carried out using glutaraldehyde, and the permeability of the shell could be controlled by the extent of its reaction. The effect of restricted shell permeability was seen through prednisone acetate release, where the cumulative release from 50 and 100% crosslinked micelles was 17.1 and 22.8% after 170 h. Increasing the temperature to 38 \({}^{\circ}\)C, above the LCST of PNIPAM, led to cumulative releases of 44.7 and 51.2% over the same time period, showing an accelerated yet still very sustained release. (BA)(AC)\({}_{2}\) (A = PMMA, B = PPEGMA, C = PNIPAM) miktoarm polymers were synthesized through a combination of CuAAC coupling with ATRP, on a 1-azido-2,3-propanediol core (Figure 12) [65]. These miktoarm stars contained only hydrophobic blocks emanating from the core junction, with hydrophilic chains being linked to hydrophobic PMMA chain-ends. While it was speculated that this structure may result in especially low CMCs, it was measured at 2 mg/L, in line with most other miktoarm star polymers. An interesting aspect of this miktoarm star architecture stemmed from the inclusion of PNIPAM arms, which, upon micellar self-assembly at room temperature, were found to be in the coronae of micelles alongside PPEGMA, but collapsed into the PMMA cores at physiological temperatures. When the temperature was further increased above the LCST at 42 \({}^{\circ}\)C, it resulted in micellar aggregation.