Updates on polyurethane and its multifunctional applications in biomedical engineering

Thermoplastic polyurethane (TPU) is a type of PU that can be melted and reshaped multiple times. They have good abrasion resistance, high elasticity, and good chemical resistance. Polyols include di-hydroxyl terminated macroglycols of polyethers, polyesters, and polycarbonates between 1000 and 5000 Da molecular weight. The molecular weight and polyol's structure are significant in PU's mechanical, physical, and chemical properties [26, 27]. TPUs are non-crosslinked polymers typically made by reacting a polyisocyanate with a polyester or polyether polyol, chain extender, and catalyst. The resulting polymer combines hard and soft segments, with the hard segments giving TPU its strength and stiffness and the soft segments giving it flexibility and elasticity. The properties of TPU can be tailored by varying the ratio of hard segments to soft segments and using different polyisocyanates and polyalcohols (figure 2) [28]. These components define the polymer structure. The soft segment determines the elasticity and flexibility of the TPU and is typically a polyetherdiol in light of its hydrolytic stability in biomedical applications [29]. The hard segments consist of an aromatic di-isocyanate, such as MDI, which gives the polymer its thermoplastic attributes.

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A low-molecular-weight diol is also used as a chain-extending agent [28]. One of the key features of TPU is its thermoplastic nature, which allows it to be melt-processed and molded in various shapes and forms. When the ratio of soft and hard segments is high, the PU is more deformable and exhibits a rubber-like behaviour. The soft segments in the polymer structure provide flexibility and resilience, allowing the polymer to stretch and recover its original shape without permanent deformation. When the ratio of soft and hard segments is low, the PU is less deformable and exhibits a more complex and rigid behaviour.

The hard segments in the polymer structure provide strength and rigidity, limiting the polymer's ability to stretch. It is worth noting that the ratio of soft segments to hard segments can be adjusted to achieve the desired balance of properties [31, 32]. For example, a PU with an average ratio of soft segments to hard segments can exhibit a balance of flexibility and strength, making it suitable for various applications.

The polymerization of TPU uses aromatic bifunctional reagents diisocyanate (MDI) and a high-molecular-weight polyether-diol. The short chain-extending diol increases the polymer's molecular weight; e.g. Texin 986 uses MDI and a high-molecular-weight polyether-diol [33]. A short chain-extending diol increases the hard segment and the molecular weight of the TPU [34, 35]. In general, thermoplastic polyether-urethanes, which comprise the most implantable materials, have very high tensile strength, toughness, abrasion resistance, and resistance to degradation, in addition to biocompatibility, which has sustained their use as biomaterials [36]. The stability of thermoplastic polyether-urethanes depends on the polyol groups used in the synthesis [37]. TPUs are suitable for medical devices such as artificial joints, heart valves, and blood vessels. They are also used in coatings for medical implants, such as stents and artificial heart pumps, to reduce thrombosis [38, 39]. TPU can also be used in wound dressings, sutures, and TE applications [40]. Poly(ether urethanes) are used in blood-contacting applications, such as catheters, heart assist pumps and chambers for artificial hearts, vascular prostheses, drug delivery, and pacemaker lead insulation [24, 41, 42].

Poly(ether urethanes) are made from polyether polyols and are known for their excellent resistance to hydrolysis and good flexibility at low temperatures [43].

Poly(ester urethanes) are made from polyester polyols and are known for their good abrasion resistance, high tear strength, and excellent chemical resistance to oils and solvents. However, it is unstable towards enzymes and has a high degradation rate. Therefore, they can be utilized as bioabsorbable polymers [43].

Poly(carbonate urethanes) are made from a combination of polycarbonate and polyester or polyether polyols, known for their high strength, high temperature, and good chemical resistance [44, 45]. PUs are used due to their physiological compatibility, appropriate haemocompatibility, excellent in vivo stability over long implant periods, and excellent physical and mechanical properties [8, 46]. Polycarbonate urethanes have displayed significant potential as durable elastomers that remain stable over extended periods. They demonstrate exceptional resistance to processes such as hydrolysis, environmental stress cracking, and oxidation caused by metal ions [47]. Different polyol groups can affect the properties of the resulting material, such as its flexibility, toughness, and thermal stability. The choice of polyol groups can also affect the material's processing conditions and final properties [26, 48]. Table 2 shows the biomedical applications of various TPUs.

It has been shown that polycarbonate-based urethanes have better biostability and biocompatibility than other types of PU (e.g. polyester urethane) for biomedical applications, which were correlated to better and higher stability of polycarbonate's soft segments inside of the body [56]. Zhu et al have designed a system containing 4,4'-methylenebis (cyclohexyl isocyanate) (H12MDI), poly(1,6-hexanediol) carbonate diols (PCDL), 1,4-butanediol (BDO) and 1,6-hexamethylene diisocyanate (HDI) to synthesize polycarbonate-based urethanes and evaluated biocompatibility property using platelet adhesion and haemolytic tests. They reported that this composition could be applied as a biomaterial due to its excellent blood biocompatibility [57]. Previous reports showed that polycarbonate-based urethanes have higher biostability and oxidative stability than polyether urethanes, which caused them to be a better replacement for polyether urethanes for biomedical applications [29, 56]. They showed appropriate heat stability, mechanical properties, and hydrolytic stability; however, they are reported to undergo possible enzymatic degradation because of inflammatory cells for long-term in vivo investigations [58].

Biocompatibility refers to the ability of a material to perform with an appropriate host response in a specific in vivo application. In other words, it means that a material does not cause any adverse non-physiological reaction when it comes into contact with living tissue or bodily fluids. As with all biomaterials, PU's biocompatibility can vary depending on the specific application and the type of tissue or fluid with which the material will come into contact. Therefore, biocompatible materials in one application may not be biocompatible in another one [59]. However, cytotoxicity is the minimum requirement for using a PU for medical applications [60]. PU's cytotoxicity has been extensively investigated, including in vivo and in vivo studies. For example, PU elastomer was previously evaluated for its blood biocompatibility [61, 62]. In another case study, mouse fibroblast cell lines and human saphenous-vein endothelial cells (HSVECs) were used for culture with various commercial PUs. The results showed that PU could provide sufficient conditions to colonize patient-derived HSVECs [63]. Since different PUs have been developed, it is vital to understand the relationship between biocompatibility and PU structure. Lyman and Picha reported that the biocompatibility of PUs depends on their configuration and morphology [64, 65].

Moreover, the relationship between soft and hard segments and the blood response has been reported in other investigations. They illustrated that the hard segment is widely thrombogenic [66, 67]. Two reasons were considered for this high thrombogenicity; the first was the high-level crystallinity of polymers, and the second one was that they demonstrate a strong and reasonable hydrogen bond surface. Cooper et al stated that surface mobility could be essential in interacting with polymers and biological systems [68]. In the late 1980s, many studies investigated chemical, structural and morphological modification in PUs to confirm their blood biocompatibility. Recently, some surface modifications and synthesis methods, such as grafting, chemical incorporation, and coating techniques, have been developed to increase the blood biocompatibility of PUs [69]. Incorporating and substituting nanoparticles is another way to improve PU biocompatibility [70–72]. In addition to the composition variation, ion beams and plasma techniques have been considered an appropriate way to enhance PU biocompatibility [73].

Biostability refers to the ability of a material to resist degradation by microorganisms, such as bacteria and fungi. Biodegradability, on the other hand, refers to the ability of a material to be broken down by microorganisms over time. Some PU formulations can be biostable, meaning they are generally resistant to degradation by microorganisms (table 3). Certain PUs can be made biodegradable by incorporating biodegradable polyols into the polymer structure. These biodegradable PUs are typically made from polyester polyols modified to include groups that microorganisms can easily break down [74].

The degradation mechanism of biodegradable PUs typically involves the action of enzymes produced by microorganisms. The enzymatic degradation of PU is shown in figure 3. The enzymes break down the PU polymer by attacking the ester bonds in the polyol, causing the polymer to lose its mechanical properties and eventually break down into smaller molecules that microorganisms can further degrade [75]. Figure 4 illustrates PU biodegradation.

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It is worth noting that the biodegradation rate of PU polymers is affected by various factors, such as the type of polyol used, the environmental conditions (e.g. temperature, humidity, and the presence of other microorganisms), and the size and shape of the polymer. Therefore, the biodegradability of PU polymers is highly dependent on the specific conditions of the environment. PU can degrade through various mechanisms depending on the type of polyol used. Degradation mechanisms can co-occur, and their degradation rate is affected by environmental conditions. Therefore, choosing the appropriate PU type is essential depending on the intended application and environmental exposure conditions [10]. Some of the possible degradation mechanisms include:

Ester hydrolysis: PUs made from polyester polyols are vulnerable to hydrolysis, a chemical reaction that occurs when the water breaks down the ester bond in the polyol. Hydrolysis can cause the PU to weaken and lose its mechanical properties over time [10].

Ether oxidation: PUs made from polyether polyols are vulnerable to oxidation, which is a chemical reaction that occurs when the ether bond in the polyol is broken down by oxygen. This can cause the PU to become brittle and lose flexibility over time [76].

Heat-induced oxidation (HIO): is a degradation mechanism that can occur in polyurethanes (PUs) when exposed to high temperatures. HIO can cause the PU to become brittle and lose its mechanical properties, as well as a change in colour and surface appearance. During HIO, the polymeric chains in the PU can break down due to the presence of free radicals generated by the high temperature. This can lead to the formation of short-chain fragments, which can cause crosslinking between the polymer chains. While this crosslinking can increase the PU's thermal stability, it can also cause the material to become brittle and lose its mechanical properties. The brittle behaviour of PUs at high temperatures instead of melting is due to the nature of the polymer structure. PUs typically comprise soft and hard segments, resulting in a complex morphology that can affect the material's behaviour at high temperatures. The hard segment in the PU can act as a stabilizing agent, preventing the material from melting even at high temperatures, but when exposed to HIO, the hard segments can also degrade, causing the material to become brittle. While increasing the temperature can cause crosslinking in PUs, it can also lead to HIO, which can cause the material to become brittle and lose its mechanical properties. Understanding the underlying mechanisms of PU degradation at high temperatures is essential for developing PUs with improved thermal stability and better resistance to HIO [77].

Environmental stress cracking (ESC): is a phenomenon that can occur in PUs when exposed to certain chemicals and mechanical stress. ESC can cause PU to become brittle and lose its mechanical properties, resulting in cracks and failure of the implanted device. ESC is more commonly observed in polyether-based PUs (PEUs) than polyester-based PUs (PESs) due to ether linkages in the polymeric chain. When exposed to certain chemicals, these ether linkages are more susceptible to hydrolysis, forming cracks and fissures in the material. Mechanical stress, such as bending or flexing of the material, can further exacerbate ESC's effects by increasing the crack propagation rate. The synergistic effect of chemical exposure and mechanical stress can cause the PU to fail, leading to potential complications in medical device applications. ESC is a significant concern, as PUs are commonly used in implanted devices such as catheters, stents, and artificial heart valves. In these applications, the PU may be exposed to various chemicals in the body, such as blood, enzymes, and medications, which can lead to ESC. In conclusion, ESC is a phenomenon that can occur in PUs, especially in polyether-based PUs when exposed to certain chemicals and mechanical stress. Understanding the underlying mechanisms of ESC is essential for developing PUs with improved chemical and mechanical stress resistance, which can increase the safety and efficacy of implanted medical devices [78].

Enzymatic degradation: PUs are known to be biodegraded by certain microorganisms, particularly bacteria and fungi. Enzymes can cause PU to become weaker and lose its mechanical properties over time [76, 79, 80].

Carbonate degradation: PUs made from polycarbonate polyols are vulnerable to degradation caused by the presence of the carbonate group in the polyol, leading to a loss of strength and mechanical properties [10, 81, 82].

Biostability is another essential property for the long-term medical application of PU [36]. Polyester-based PU is known not to be resistant to hydrolysis. The ester-group starts to degrade after implantation, leading to strong inflammatory reactions. This problem was solved by introducing a hydrolysis-stable polyether-urethane [83] based on polyether. Hence, polyester-based PUs are no longer used to produce implants [36]. Generally, the hydrolytic degradation of polyether-urethanes in pure water is minimal. Nevertheless, the aqueous environment of the body, i.e. cations and anions, has a strong catalytic effect in hydrolysis. Polyether-urethanes and polycarbonate-urethanes are generally less susceptible to hydrolysis than polyester-based ones. However, they can degrade hydrolytically at high temperatures (>50 °C) in the presence of water. These conditions often exist in polymer processing methods like injection molding and extrusion [84] and should be avoided.

For this reason, most PUs available for medical applications are based on polycarbonate or polyether. However, the biodegradation property of PU has been widely investigated [8, 85, 86]. In this regard, some modifications have been reported to improve PU biostability.

One proposed modification is to replace the soft domains in the PU with materials that have improved stability, such as polysiloxane or polyolefins [87–89] since it has been reported that the soft segments had the most prominent effect on biostability [90]. Haugen et al found that polyether-urethane scaffold biostability was worsened with a higher gamma-irradiation dose [79, 91]. Cozzens et al [88] designed a new PU based on a mix of poly(tetramethylene oxide) (PTMO)/poly-isobutylene (PIB) and provided accelerated testing (H₂O₂ (20%) solution contained 0.1 (M) CoCl₂ at 50 °C) to evaluate and analyse resistance against degradation of metal ions oxidative in vivo. The new composition showed considerable stability compared to the commercial types, like Pellethane 2363-80A and 2363-55D [88].

Additionally, surface modification [92], antioxidants [93], and nanoparticles [94, 95] were also introduced as methods for the improvement of PU stability. Notably, the biostability property is vital for long-term applications since biodegradation entails a loss of shape and mechanical properties. However, investigation of some scaffolds-based biodegradation in PU is also needed in TE, and control of the degradation rate is essential in some applications, such as drug delivery, scaffolds, and short-term implants [96, 97]. Many factors can affect PU's biostability, including mechanical properties, chemical structure, morphology, fabrication methods, and working condition [95]. Surface chemistry, the relative hydrophobicity, and hydrophilicity of polyether-urethanes have been found to influence the proteins adsorbed by the surface [37]. The first event in soft tissue interactions with polyether-urethane is the adsorption of water and protein onto the surface, followed by an acute inflammatory response characterized by polymorphonuclear leukocytes. The consequence of the adsorbed protein layer on the biomaterial is that cells never actually contact the material. After protein adsorption, an acute response is followed by chronic inflammation, which includes a wound-healing mechanism and foreign body reactions involving macrophages and fibroblasts [59]. Like many other biomaterials, polyether-urethanes elicit a foreign body response determined by multinucleated giant cells in contact with their surfaces, generally called ESC.

Additionally, researchers have found that polyether-urethanes often become encapsulated in a fibrous capsule [36, 99]. Ultimately, this influences the subsequent cell migration and implant function [100]. Mohanty et al [101] and other researchers [102] have found a relationship between pore size and the infiltration of macrophages into a polymer implant. The researchers found that a 5–10 µm pore inhibits cell infiltration in the bulk material, while larger pore sizes (200–300 µm) encourage cell infiltration [36]. Mohnty's work also indicates that a porous implant delays the incidence of capsule formation and contracture for about ten years. Haugen et al reported using porous PU as a gastro application to reduce gastroesophageal reflux diseases. Where Haugen et al developed a porous PU implant with supercritical CO₂ [103, 104], where the porous structure was supposed to fixate around the oesophagus. Recently, in vivo experiments have shown that tissue ingrowth into a porous PU structure is only possible with plasma surface treatment, and untreated TPU resulted in massive fibrotic encapsulation [105].

Hence, chemical composition is the key factor [86]. ESC is also reported as one reason for PU's degradation [106]. Various works have investigated other mechanisms of PU degradation [5, 9, 97, 107].

In summary, over the last two decades, several chemistry approaches have been designed primarily focusing on decreasing or eliminating functional groups sensitive to oxidative degradation to improve biostable PUs. in vivo analysing and animal investigations have produced long-term biostability for clinical applications. While most approaches have helped design PU elastomers with improved resistance to oxidative degradation compared with conventional PU, their utility in long-term medical implants is yet to be assessed. The improved understanding of the relationship between PU chemical structure on properties and biostability will help synthetic polymer chemists design new materials to suit the needs of next-generation medical implants.

Smart materials known as shape memory polymers (SMPs) can remarkably alter and regain their shape in response to various stimuli [112]. The shape memory of polymers has been investigated previously [113]. Shape memory polyurethane (SMPU) is a TPU that exhibits shape memory properties. When SMPUs are heated above their glass transition temperature (T_g) but below their melting temperature (T_m), the material can be deformed into a temporary shape maintained by cooling it below the T_g. Soft segments in the PU structure are responsible for this temporary shape, as they allow the material to be deformed and maintain the temporary shape. Upon heating the material above the T_g, the hard segments in the PU structure help to recover the material's permanent shape. The hard segments are responsible for 'remembering' the permanent shape and help drive recovery [114, 115]. Tm is an essential parameter for processing SMPUs to achieve the desired permanent shape. During processing, the material is typically heated to a temperature above the T_m to facilitate the recovery of the permanent shape. The shape memory behaviour is achieved by adjusting the polymer structure's ratio between hard and soft segments. The hard segments provide the polymer with shape memory properties, while the soft segments provide flexibility and elasticity [116]. SMPU can be a good candidate for medical devices because of their flexibility in shape change in different areas [117, 118]. A specific chemical combination can allow a geometric shape-changing that makes them resistant to external factors like pH [119], light [120], and temperature [121], as shown in figure 5. This change is because the polymer's structure contains switching segments [122]. If the temperature exceeds the glass transition temperature (T_g), materials are in their rubbery-elastic form, which can deform easily in the desired shape. If the temperature falls below T_g, they become fixed without any deformation in shape. At this step, materials are rigid. For biomedical applications, T_g should be around the body temperature to allow the recovery of the permanent shape in vivo. This phenomenon occurs due to molecules' movement and structure [117]. This shape memory property can make these polymers an ideal material for medical devices because they can keep a specific shape during implantation or delivery and recover to their original shape. Interestingly, increasing the temperature above T_g can recover the materials' original, permanent shape. The development of SMPUs began 15 years ago [117]. The biomedical applications of PUs as SMPs include clot removal, vascular occlusion, and stenting [123, 124]. Also, it has been reported that SMPs can be used for cardiovascular implants [53], bone tissue engineering (BTE) applications [116], electromagnetic shielding, pressure bandages, and self-healing [125]. Also, range of dual bioactive electroactive shape memory PU elastomers by incorporating dopamine into a formulation has been synthesized based on citric acid, PCL, dopamine, and the electroactive aniline hexamer. These materials are suitable for soft tissue applications that require sensitivity to electrical like skeletal muscle regeneration [126]. A pH-sensitive PU was also designed.

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In this case, the shape memory of PU depended on pH variation instead of temperature variation. The mentioned PU was designed for drug delivery and TE applications [127–131]. It is worth noting that the shape-memory effect is highly dependent on the specific conditions of the environment, such as the temperature, humidity, and the type of polyol used [132]. De Nardo et al showed that SMP scaffolds could be obtained via solvent casting/particulate leaching of gelatin microspheres prepared via oil/water emulsion. Varying the gelatin microsphere size enabled easy control of scaffold morphology, pore size, and shape. Homogeneous spherical and interconnected pores have been achieved with the preservation of shape memory ability, with a recovery rate of up to 90%. Regardless of pore dimensions, MG63osteoblast-like cells were observed adhering and spreading onto the inner surface of the scaffolds obtained in in vivotests [118, 133].

In summary, using SMPUs in various applications demonstrates their versatility and importance. The shape memory effect can be driven by external forces such as electricity, temperature, electricity, and light. SMPs respond to external stimuli because of hard and soft segment domains. Desired properties of SMPUs can be achieved by cross-linking through ionic, covalent, and hydrogen bonding [128].

Some types of PU have been designed to exhibit antibacterial properties. This can be achieved by incorporating antimicrobial agents, such as silver or zinc ions, into the polymer structure. These antimicrobial agents can kill or inhibit the growth of bacteria, fungi, and other microorganisms on the surface of PU devices [136]. The antimicrobial agents are incorporated into the polymer structure during the manufacturing process, and they are released slowly over time, providing long-lasting protection against microorganisms. The antimicrobial agent's release rate can be controlled by adjusting the concentration of the antimicrobial agents in the polymer and the molecular weight of the polymer [137]. Antibacterial PU can be used in many applications where controlling microorganisms' interaction with the PU structure is essential, such as in medical devices and food packaging. Like any other antimicrobial agent, the effectiveness of the antibacterial properties of PU can be influenced by many factors, such as the amount of antimicrobial agent used, the type of microorganism, and the environmental conditions [138]. Therefore, choosing the appropriate type of PU and antimicrobial agent is essential, depending on the intended application and the microorganisms needing control [139].

Kasi et al comprehensively collected these investigations. Different methods and materials have been reported to increase antibacterial properties [140]. Moreover, silver (Ag) nanoparticles, metal oxides and metals, nanomaterials based on carbon, synthetic polymeric materials based on PU, natural sources, and more have been reported [140]. In a different study, functionalized PU was applied as a magnesium (Mg) coating to enhance and improve antibacterial and corrosion resistance. The results showed considerable enhancement in corrosion resistance and excellent antibacterial activity [141]. Jiang et al designed amphiphilic poly(dimethylsiloxane) based on PUs, using carboxybetaine to obtain an antibacterial effectiveness of 97.7% [142]. Because PUs have remarkable properties, they have been used as a based substrate for other antibacterial agents, such as copper oxide (CuO) and chitosan [143, 144]. The effects of different agents, such as Ag, zinc oxide (ZnO) [145], and titanium dioxide (TiO₂) nanoparticles loaded in a PU matrix, have been reviewed, and it has been reported that these nanoparticles can improve PUs' antibacterial activity [146]. Wang et al reviewed the antimicrobial PU coatings for different applications [147]. Also, Wang et al evaluated the antibacterial activity of PU and suggested different methods to design PUs' antibacterial properties for use in medical devices [136].

PUs have been investigated as a material for drug delivery systems due to their unique features and have been used as PU-based hydrogels, PU-based nanoparticles, PU-based microparticles, and PU-based films and membranes [203]. For example, they can be used as a coating on implantable devices to control the release of drugs. The properties of PU can be modified to suit the specific drug delivery application by adjusting the polymer properties, such as porosity, mechanical properties, and degradation rate [204]. Controlled drug delivery has been an important issue in recent years because of its advantages, such as lower toxicity and less time required for injection [205, 206]. Also, selecting a carrier for the drug is critical because it can protect drugs from damage or loss or play an essential role in providing a targeted and smart release [207]. Carriers are applied to immobilize the active materials and substances in drug delivery preparation. They should have a balanced structure between hydrophobicity and hydrophilicity and good biocompatibility [208, 209]. The form of a carrier can vary based on the targeted application [210]. PUs are an attractive material for drug carriers and drug delivery systems because of many possible modifications and processing techniques that can be applied to control drug release systems. PU nanoparticles, scaffolds with various porosity and pore sizes or coatings, different compositions and hydrophilic properties of raw materials, type, and the ratio of hard and soft segments, crystallinity, and the crosslinking degree have the properties and forms that give the possibility to control speed and amount of drug released. Also, modified cross-linked PUs can behave as hydrogels, absorbing large amounts of water without dissolving, an essential quality for drug carriers [207].

Depending on the targeted application of PUs, they can be classified into two groups: biodegradable and bioinert. The first category is used in drug delivery [211]. PU scaffolds or stents have been recognized as a suitable carrier for drug delivery, and many investigations have been conducted. Moon et al loaded heparin-deoxycholic acid (DOCA) into a PU coating and reported that the percentage of loaded DOCA on the PU was more than 75% and that the released DOCA could prevent fibrin clot formation or adhesion of platelets on the film and layer surface [212]. Moura et al demonstrated the potential for drug delivery using PU-based implants by loading dexamethasone onto the material. Their findings showed that using PU for drug delivery could modulate the fibrosis and angiogenesis induced by the disc-shaped spongy and main inflammation components [213]. Other drugs in molecule shape have been investigated, such as 5-flurouracil and ibuprofen [214–216]. In addition to the molecular-shape drugs, rhBMP-2 is a protein based on recombinant human bone morphogenetics incorporated into scaffold-based PU and implanted for defects healing in the rat femoral bone. The results showed that the formation of bone increased after four weeks of implantation [217].

In another study, a PU that contained 5-aminosaheylic acid (5-ASA) was used, and results showed that the urethane group hydrolysis caused a release of drugs [218]. Also, Ghosh and Mandal reported on the release of ibuprofen based on PU structure [219]. It has been reported that drug incorporation does not considerably affect PU's biological and mechanical properties. Simmons et al reported that after using dexamethasone acetate in PU based on siloxane, there were no considerable differences in biocompatibility and biostability after six months of implantation in an animal model [220]. Drug delivery by PU-based carriers can be affected by some factors. For example, Da Silva et al investigated that the presence of poly(caprolactone) together with poly(ethylene glycol) as a soft segment in PU could enhance the release rate of dexamethasone acetate compared to PU with poly(caprolactone) alone. Furthermore, it was shown that the structure of PU was also involved in drug release behavior [221].

Furthermore, the use of ionic ligands has been shown to accelerate the release of drugs from PU-based carriers [222]. It has also been reported that the drugs' properties and functions can influence the drug release rate from the PU carrier [223]. This suggests that the drug itself plays an essential role in determining the release kinetics of the carrier. The factors related to the drugs' properties and functions, such as their molecular weight, hydrophobicity, and charge, can influence their interaction with the PU carrier, thereby affecting the drug release rate. For example, drugs with a higher molecular weight or greater hydrophobicity may have a slower release rate from the PU carrier, as they may be more strongly bound to the material.

Similarly, drugs with a higher charge may interact more strongly with ionic ligands, resulting in a faster release rate. Moreover, using ionic ligands can accelerate the release of drugs from a carrier base. E.g. Grigoreva reported that changing the PU structure (soft and hard segments) could make it possible to design composite PUs based on desired drug delivery. Another example; is the release of drugs, such as naltrexone, cefazolin, and piroxicam, on the base cross-linked PUs and dioxydine for linear PUs was shown to have an effect [224].

Gentile et al investigated a system based on ceramics scaffolds coated with gelatin. Then, they incorporated PU nanoparticles into the gelatin, loaded with indomethacin (IDMC), into the scaffold. Incorporated PU nanoparticles allowed a sustained IDMC release at about 65%–70% during the first week, and compressive modulus was increased [225]. Kolmas et al also used a system containing PU and HA to avoid cell invasion and the growth of bone tumours in the extracellular matrix (ECM). The system was designed to facilitate a prolonged release of bisphosphonates. The results showed that the release rate varied from 20% to 80%, indicating a possibility of controlling the release kinetics [226]. Guo et al worked on stent coatings based on biodegradable PU and were able to adjust the drug release [21, 227]. Polymer nanoparticles were analysed for improvement in cancer chemotherapy. Therefore, declining drug resistance and improving synergy in malignant lesions are necessary. Nanoparticles based on PU can co-encapsulate chemotherapeutic agents, such as doxocetal and doxorubicin hydrochloride [228]. Tissue adhesives based on PU can also be applied as a delivery system to release antibiotics and painkillers [171] slowly. Some types of PU that have been investigated so far are summarised in table 4 [229].

In conclusion, PUs family is widely used in biomedical applications. According to the adjustable physical and chemical properties of PUs, it is possible to change the degree of crystallinity and their interaction with water. Since PUs have appropriate biocompatibility, low cytotoxicity, and good mechanical properties, they have the potential to overcome challenges in drug delivery systems [218, 219]. By further attempting to have more in vivo investigation, more applications, limitations, and challenges of PUs for controlled drug delivery can be found.

PU has been investigated as a biomaterial for various tissue TE applications due to its unique properties, such as biocompatibility, flexibility, strength, and durability. Scaffolds are critical to TE as they support cells to grow and differentiate into functional tissue. The properties of PU, such as biocompatibility, flexibility, and durability, make it an attractive material for scaffolds [244]. PU-based scaffolds can mimic the mechanical properties of natural tissue and can be designed to provide a suitable environment for cell growth and differentiation [245]. PU-based scaffolds have been developed for various TE applications, including skin, cartilage, bone, and nerve tissue. For example, PU-based scaffolds have been developed to repair skin for treating burns and wounds. These scaffolds can support cells to grow and differentiate into functional skin tissue [246]. PU-based scaffolds can also be used to repair cartilage and bone tissue. These scaffolds can support cells to grow and differentiate into functional cartilage and bone tissue [247]. PU-based scaffolds have also been used in nerve TE, as they can mimic the mechanical properties of the nerve and provide a suitable environment for the growth and differentiation of neural cells [248]. PU scaffolds can have a variety of morphologies depending on the manufacturing process and the specific application [60]. The morphology of a PU scaffold refers to the physical structure and shape of the scaffold. PU can be fabricated as a porous structure (figure 10), to improve and enhance proliferation, cell adhesion, and differentiation.

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Some typical morphologies of PU scaffolds include:

• (1)  
  Foam: PU scaffolds can be made in foam, a porous, spongy material with a high surface area-to-volume ratio. Foam scaffolds are lightweight and easy to handle, and they can provide a suitable environment for cell growth and differentiation [247, 250].
• (2)  
  Fibrous: PU scaffolds can also be processed in thin fibres and elongated structures that mimic natural tissue structures. Fibrous PU scaffolds can provide a suitable environment for cell growth and differentiation. They can be used in many TE applications, such as tubular structures for blood vessel regeneration, nerve guides for peripheral nerve regeneration, and cartilage TE constructs [248, 251].
• (3)  
  Film/coating: PU scaffolds can also be made as a film. Film scaffolds can be used in many applications, such as wound dressings, cell culture substrates, or coating materials to modify the substrate's surface properties to provide an active environment for cell growth and differentiation or an additional layer of protection [141, 252–255].

Figure 11 shows a hybrid gelatin-based electrospun scaffold that the use of biodegradable polycarbonate PU has fabricated. It is supposed that adding PU could enable a tailored degradation rate and an increased mechanical strength compared to electrospun gelatin. Adding 20% PU to gelatin scaffolds (gelatin 80− polycarbonate PU 20) significantly increased the scaffolds' yield strength, degradation resistance, and elongation. Gelatin 100 had an elastic modulus of 130 ± 10 kPa, whereas the addition of 20% PU enhanced the elastic modulus to 260 ± 11 kPa. Gelatin 80−PU 20 showed the greatest elongation at break (60.7 ± 4.8), while this parameter was 34.7 ± 7.7 for gelatin 100. In vivo studies using a mouse excisional wound biopsy grafted with the scaffolds demonstrated that the gelatin 80−PU 20 scaffold enables more significant cell infiltration than clinically established matrices.

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The findings show that electrospun gelatin 80−PU 20 scaffolds can generate tissue substitutes and overcome some limitations of conventional wound care matrices [256].

Topography can control the behaviour of cells, such as proliferation and growth alignment. Moreover, the shape change (stimulus-regulated swelling and shrinking) benefits the release of encapsulated agents in a precisely controllable manner [257, 258]. Additionally, the scaffold should benefit from a targeted degradation rate, and this rate should be designed based on the new tissue formation rate. Modifying PU with ascorbic acid (vitamin C) can improve material properties for applications in soft tissue [259]. Shahrousvand et al also incorporated some iron oxide nanoparticles in PUs for TE applications [260]. Results indicated that the produced material had good potential in cell therapy applications. Li et al fabricated a scaffold containing hydroxyapatite (HA)/PU. Glyceride of castor oil (GCO) was applied to modify the soft part of PU. Finally, it was added to enhance the compressive strength, which increased when the HA particles were added.

Moreover, the HA particles were essential in bonding with the bone tissue and GCO-PU matrix interface [261]. HA and PU have been investigated more because HA can improve bioactivity, and PU can provide an excellent mechanical property [262]. Haryńska et al designed a 3D-printed filament based on PCL and PU for medical applications [14]. Also, thermoplastic-elastomer PU has been reviewed for its potential as a filament for FDM 3D printers [263]. Piotrowska-Kirschling conducted a study on the combination of chitosan and PU. It was reported that the presence of chitosan accelerates PU degradation. The combination of chitosan and PU was introduced as a highly biocompatible and bioactive combination that can be a good candidate in TE applications [264].

Additionally, the combination of chitosan and chitin with PU caused the creation of an applicable and influential composite for the biomedical field. The amino, carbamate, and acetamide groups (in chitosan, PU, and chitin, respectively) play a prominent role in making a material suitable for medical applications. In another investigation, collagen and PU were reviewed for TE application [265].

The use of PU for 3D-printed scaffolds is currently receiving much attention. The methods studied are fused filament fabrication (FFF), inkjet printing, stereolithography (SLA), and low-temperature deposition. Scaffolds based on poly(carbonate-urea) urethane enhance surface adherence and cell viability; therefore, they can become a polymer to improve laryngeal regeneration and reconstruction [266, 267]. In addition, elastomeric-PU can avoid shear forces between implant and bone, supporting osteogenic cells' proliferation in hard tissue. Regarding soft tissue, PU-based scaffolds can be applied to muscle regeneration, blood vessels, heart tissue, cartilage, and nerve [268].

A novel PU scaffold containing megni oil was used in TE due to its ability to enhance antithrombogenicity [269]. On the other hand, angiogenesis is another hot topic in the biomaterials field, and many researchers are trying to design a material that either inhibits or improves it. Colloidal aggregates and gels based on the mediated aggregation of cationic-PU particles have been applied to produce a matrix [270–272]. Notably, PU-based scaffolds have been established for the repair and regeneration of bone tissue based on osteogenic differentiation and proliferation [52, 273]. Membranes of PU and silk fibroin were designed and evaluated to mimic oral soft tissue [274]. The biological effect of these membranes was analysed regarding cell proliferation, cell adhesion, phosphatase activity, calcium content, and viability. These membranes seemed adequate for bone tissue regeneration [274, 275].

It is worth noting that the scaffold properties can be modified to suit the specific TE application by adjusting the polymer properties, such as porosity, mechanical properties, and degradation rate. Moreover, developing these scaffolds is still ongoing as the field of TE is rapidly advancing.

Thoughsome major challenges persist, the future of PU scaffolds is promising. The PU scaffolds in a specific form targeted for a specific organ or tissue can be investigated. Thus, for future recommendations, a multidisciplinary approach is still required for fully exploiting PU scaffolds, combined with computational approaches for designing patient-specific PUs TE scaffolds. As a result, more investigations are required to narrow the differences between the cell interactions of PUs scaffolds in the in vivo and in vivo environments [276].

The wound-healing process involves regenerating epidermal and dermal tissues that heal the wound. Successful wound healing involves interaction between dermal and epidermal cells, angiogenesis, and ECM. These can be regulated by growth factors and cytokines [277]. It has been reported that an ideal wound dressing should have specific characteristics and properties, such as a high level of bacterial resistance to removing exogenous microorganisms, sufficient haemostatic properties, high fluid adsorption, specific mechanical properties, stability, the ability to be removed easily, cost-effectiveness and recyclability [278]. In this regard, Unnithan et al investigated the fabrication of nanofibre composites for wound dressing based on cellulose, PU, and zein with streptomycin sulfate. The results showed perfect cell attachment, antibacterial activity, hydrophilicity, and blood clotting [137]. Also, mats based on nanofibrous composites containing gelatine, chitosan, and SMPU were fabricated by electrospinning and post-treated with silver nitrate (AgNO₃) solution. The obtained data illustrated that the produced composite could be a good candidate for wound dressing because it showed adequate antibacterial activity, an appropriate water vapour transmission rate, surface wettability, cytocompatibility, and haemostatic properties [279]. In another investigation, a combination of hydrogel and PU produced a composite for wound healing. The fabricated composite could accelerate wound healing and reduce the formation of scars [280]. PU compositions are selected for wound healing applications based on their chemical composition, shape, and functionalization [7, 281]. Transparent and thin PU films are primarily suitable for surface protection and wounds after an operation with less or no exudate secretion. These PUs are not permeable to water and bacteria but are permeable to gases. As a result, they can help wounds to breathe.

PU foams are absorbent and suitable for wounds with more exudate [281]. Salehi et al designed a nanofibre matrix based on polyurethane and hyaluronic acid (PU-HA). This nanofibre matrix was enriched with various concentrations of ethanolic extract of propolis (EEP) (figure 12, top panel).

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Results demonstrated that the PU-HA/1% EEP structure could be a good candidate for wound healing, with promising biocompatibility and antibacterial activity. It has been reported that PU in the composition accelerated the wound healing process (figure 12, lower panel) [282]. Also, lignin-based polyurethane foams loaded with Ag nanoparticles were investigated as a composite that showed good antibacterial activity in wound healing applications [283]. Even incorporating copper (Cu) in PU composites was reported to fabricate the nanofibrous composite for wound healing [284]. Another composition including PU for wound healing application has been introduced by Li et al [285]. A dressing material has been developed made of PU that possesses multiple functionalities. The designed material combines the electroactivity of aniline oligomer segments, the wettability of PEG segments, and the mechanical properties of PCL segments. The films exhibited favourable hydrophilic properties and swelling ratio, exceptional mechanical strength and shape memory capabilities, electroactivity, effective free radical scavenging capacity, non-adherent properties, and biocompatibility.

PUs have perfect properties for applications as pressure sensitive-adhesives (PSAs). These materials can provide appropriate adhesion to the base substrate, such as skin, with light pressure; as a result, they can be used in surgical dressing applications [286]. These materials' adhesion strength and breathability were improved by applying various crosslinkers [286]. Dressing foams based on PUs have good absorption of water and mechanical properties but less bioactivity and poor capability for healing. So, it has been suggested that incorporating and using bioactive nanoparticles like silica within the sol-gel process can improve the rate of wound closure and accelerate the regeneration of elastin fibre and collagen [287].

In summary, PUs as a wound dressing material have been commercially in the spotlight because of main characteristics such as biocompatibility, mechanical properties, elasticity, flexibility, and good oxygen/carbon dioxide permeability [288]. PUs can be a promising candidate for wound healing applications in the future. In addition, they expanded studies on the PUs to identify their limitations, challenges, advantages, and disadvantages for wound healing applications.

PUs have been explored as a material for biosensors, devices that can detect and measure specific biomolecules such as glucose, proteins, and other biomolecules [289]. One of the key advantages of using PU in biosensors is its biocompatibility [290]. PU has good chemical stability and can withstand a wide range of pH and temperature conditions, making it a good biosensor choice [291, 292]. PU-based biosensors have been developed for various applications, including monitoring glucose levels in diabetes patients and detecting chemicals, toxins, and environmental pollutants. For example, PU-based glucose biosensors have been developed by immobilizing glucose oxidase enzymes onto a PU matrix, which can then be used to measure glucose levels in the blood [293].

Figure 13 shows a biosensor based on a nanofibre-based polydiacetylene (PDA) biosensor prepared with TPU to detect E. coli. 10,12-Pentacosadiynoic acid (PCDA) was used as a monomer for preparing PDA. In the electrospinning of PCDA and PU mixture solution, PCDA monomers were randomly distributed in the PU solution. PU-PDA nanofibre biosensors were prepared via electrospinning and tested with E. coli. The colour change in PU-PDA nanofibres began immediately when exposed to the E. coli concentration of approximately 9 × 10⁸ CFU ml⁻¹. The nanofibre mat changed colour in some spots and immediately changed to red after a few minutes [294].

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Another example is using PU-based biosensors for detecting environmental pollutants such as heavy metals, pesticides, and other toxins. These biosensors are typically based on immobilizing a specific enzyme or antibody onto a PU matrix, which can then detect the presence of a specific pollutant. A suitable biosensor should have some desired properties, including specificity, accuracy, stability, a wide range of analytical capabilities, low cost, and portability [295]. Many kinds of biosensors have been introduced to date [296], and many kinds of polymers (e.g. polyaniline, poly(vinylchloride), polypyrrole, and cellulose acetate) have been investigated as the substrate for the fabrication of biosensors [297, 298]. However, this application has widely used PU due to its appropriate biocompatibility, flexibility, cost-effectiveness, and easy processability [292, 299]. The key challenge in this route is the low solubility of hydrophobic drugs. Also, components of drugs should keep their active state until they reach the absorption site [300, 301]. In order to achieve targeted drug absorption at a specific site, polymers can be coated onto the surface of drugs or mixed and linked with the drug to prevent potential degradation [231].

While PU has many advantages and has been used in many medical applications, some drawbacks must be considered when using it in medical devices. Some of the drawbacks of the medical application of PU include [60, 163, 302, 303]:

Degradation over time: PU can degrade over time due to factors such as exposure to heat, chemicals, and microorganisms. This can lead to a loss of mechanical properties, such as strength and flexibility, and can affect the performance of the medical device [60, 163, 303, 304].

Allergic reactions: Some individuals may be allergic to PU, which can cause inflammation, itching, and other symptoms. This can be particularly problematic for long-term surgical implants [304, 305].

Incompatibility with imaging techniques: PU can interfere with specific imaging techniques, such as MRI, as it can cause image artefacts. This can make it challenging to diagnose and treat certain medical conditions accurately [306].

Sterilization issues: Sterilization problems are not unique to PU and can be challenging for many other polymers and materials used in medical devices, especially those with porous structures. Porous structures can harbour bacteria and other microorganisms, making them difficult to sterilize effectively. Furthermore, specific sterilization methods, such as gamma irradiation or ethylene oxide gas, can degrade or damage the polymer structure, reducing mechanical properties or potential toxicity. Therefore, careful consideration must be given to the choice of sterilization method and its potential effects on the polymer structure and device performance [79, 91, 307].

Cost: PU can be more expensive to manufacture than other medical device materials, such as polyethylene or polypropylene, due to the higher cost of raw materials and the more complex manufacturing processes involved.

PU requires specific chemical precursors, including isocyanates, polyols, and chain extenders, which can be more expensive than the raw materials used to produce other polymers. The production of PU also requires specialized equipment and more stringent manufacturing conditions, which can increase the overall cost of the process.

These drawbacks can vary depending on the specific type of PU used and the specific application. Therefore, it is essential to choose the appropriate PU type and consider the specific risks and benefits when using it in a medical application [302].