## 8 Functional Aspects of Honey-Loaded Scaffolds for Wound Healing
Once honey has been incorporated within the scaffold, an in-depth characterisation of the developed formulation should be conducted in order to evaluate some critical parameters to ensure that the treatment effectiveness is optimised, as summarised in Figure 2.  
The release rate of the medically active agent is one of these crucial aspects, as it determines important features such as the honey concentration at the wound and periwound area. It can be modulated to prevent the inconvenience of honey leaking from the scaffold, as shown in Figure 3, and consequently reduce the dressing change frequency, with important implications on the patients' adherence to the therapy, as the manipulation may cause discomfort [166]. Honey release is usually determined by an indirect quantificationof either sugars or markers (such as methylglyoxal for Manuka honey) liberated in the medium. The maximum absorbance at specific wavelengths can be measured on a UV-Vis spectrophotometer and plotted against time [138; 161]; alternatively, HPLC quantification can be employed to determine the cumulative release [164]. For highly hydrophilic and biodegradable polymeric structures, drug diffusion and scaffold degradation are considered the driving forces of the release process [138; 161]. On the other hand, this process of structural degradation can be counterbalanced by an increased degree of crosslinking i.e., chain entanglement, with improved rigidity and structural integrity of the network, hence improved honey retention [167]. For example, a significant reduction in methylglyoxal was observed from Ca\({}^{2+}\) and Mg\({}^{2+}\)-crosslinked gellan gum composite hydrogels when compared with the non-crosslinked equivalents [164]. Different experimental protocols are described in the literature for the evaluation of drug release in vitro, with the immersion volume being variable across the papers. For honey-alginate biopprinted scaffolds, a volume of 10 mL of phosphate-buffered saline (PBS) was chosen [165]. [PERSON] et al. carried out the in vitro release test by directly submerging the membrane in PBS buffer at a controlled temperature of 37 \({}^{\circ}\)C, but the volume of the medium was not specified [138]. Alternatively, for amorphous honey hydrogels, the sample can be inserted into a dialysis bag, as described by [PERSON] et al. [161]. A burst release is conventionally detected at the beginning of the observation time, which is followed by a plateau, in hydrogels where the drug loading was performed by the direct mixing method [161; 166]. However, a sustained release was observed from gelatine-based electrospun nanofibrous membranes, with 99 \(\pm\) 0.5% cumulative honey release only achieved after 24 h of immersion [138]. This shows that different release kinetics can be established by employing alternative tissue engineering strategies, thus allowing for a controlled extended release. The capacity of a wound dressing to swell and uptake a considerable volume when immersed in an extracellular fluid-resembling medium is indicative of its ability to absorb wound exudate and to maintain a curative moist environment on the affected skin area [168]. As with the evaluation of the drug release kinetics, the volume of the medium employed to assess the water absorption features of the scaffold is highly variable across the published literature. [PERSON] et al. conducted the swelling study by soaking a 1 \(\times\) 1 cm sample in 50 mL of PBS at 37 \({}^{\circ}\)C, and a maximum swelling of 500% of the original dry weight was recorded. However, no comparative evaluation was reported between the blank membrane and the drug-loaded ones, so a consideration could not be drawn on the contribution of honey to the swelling capacity of the unloaded gelatin nanofibrous membrane [138]. [PERSON] et al. carried out the swelling test in analogous experimental conditions (PBS, 37 \(\pm\) 0.5 \({}^{\circ}\)C) but did not specify the immersion volume.