*7.3. Bilayer Hydrogels*

Bilayer hydrogel consists of a porous architecture layer integrated with a bulk layer covalently. The porous architecture is the result of the interruption in the polymerization process. The porous layer benefits hydrogel in water retention, impact on diffusion rate, minimizing CoF and wear rate [104,119]. Gong et al. [141] developed a bilayer hydrogel with varying crosslinking degrees in the top layer. A lower degree of crosslinking resulted in high porosity and the hydrogel had a higher fluid retention capacity, which consequently minimized the CoF. The bilayer architecture formation in hydrogels is due to branch dangling chemical phenomenon [141]. A branched dangling polymer chain is achieved by polymerizing the monomers, while in contact with a hydrophobic surface. Hydrogen-rich moieties are located within close vicinity of the hydrophobic surfaces yielding a low density highly porous structure. This is attributed to the high concentration of hydrogen affecting the propagation step of polymerization. The bulk area, which is far from the hydrophobic surface, could accomplish the polymerization process due to hydrogen deficiency in this zone. Consequently, a very dense structure is formed, and the bulk area's strength enhances compared to its porous counterpart [142,143]. The SEM image of a bilayer hydrogel crosssection is presented in Figure 2. *Polymers* **2021**, *13*, x FOR PEER REVIEW 11 of 25 branch dangling chemical phenomenon [141]. A branched dangling polymer chain is achieved by polymerizing the monomers, while in contact with a hydrophobic surface. Hydrogen-rich moieties are located within close vicinity of the hydrophobic surfaces yielding a low density highly porous structure. This is attributed to the high concentration of hydrogen affecting the propagation step of polymerization. The bulk area, which is far from the hydrophobic surface, could accomplish the polymerization process due to hydrogen deficiency in this zone. Consequently, a very dense structure is formed, and the bulk area's strength enhances compared to its porous counterpart [142,143]. The SEM image of a bilayer hydrogel cross-section is presented in Figure 2.

**Figure 2.** PAAm-Alg bilayer hydrogels: cross-section view. **Figure 2.** PAAm-Alg bilayer hydrogels: cross-section view.

#### *7.4. Mechanical Testing of Articular Cartilage and Hydrogels 7.4. Mechanical Testing of Articular Cartilage and Hydrogels*

AC as a soft tissue articulates the full range of motions and experiences complex loading scenario, which is compression, tension, shear and friction [144]. Most studies focused on assessing recovered tissue based on biochemical, gene expression, or histological aspects [145]. Comprehensive protocols for mechanical evaluations showed a lack of standardization for their unit reference. Therefore, remarkable tolerances in the reported data are inevitable. The compression testing is categorized as unconfined, confined and in situ. For the confined compression test, a porous plate or indenter is used to let fluid flow out of the tissue. AC as a soft tissue articulates the full range of motions and experiences complex loading scenario, which is compression, tension, shear and friction [144]. Most studies focused on assessing recovered tissue based on biochemical, gene expression, or histological aspects [145]. Comprehensive protocols for mechanical evaluations showed a lack of standardization for their unit reference. Therefore, remarkable tolerances in the reported data are inevitable. The compression testing is categorized as unconfined, confined and in situ. For the confined compression test, a porous plate or indenter is used to let fluid flow out of the tissue.

Four test configurations are commonly used to characterize cartilage mechanical responses: ramp, stress relaxation, creep and indentation tests. Jay et al. [70] reported that the most utilized test configuration in studies from 2009 to 2018 were: ramp, stress relaxation and creep. Thus, the ramp test has been configured to simulate the load-bearing properties of the tissue. After recording the stress-strain response by the ramp test, the first-order differential equation of the curve, which is the slope of the stress-strain curve, Four test configurations are commonly used to characterize cartilage mechanical responses: ramp, stress relaxation, creep and indentation tests. Jay et al. [70] reported that the most utilized test configuration in studies from 2009 to 2018 were: ramp, stress relaxation and creep. Thus, the ramp test has been configured to simulate the load-bearing properties of the tissue. After recording the stress-strain response by the ramp test, the first-order differential equation of the curve, which is the slope of the stress-strain curve,

terials endure a higher load before ultimate failure compared to hardened materials and are suitable for replacing tissues that undergo large deformations [146]. Two factors that considerably affect the tangent modulus are strain rate and strain point. Healthy knee cartilage typically experiences average strains under 10% [36] and a maximum of 17% [147]. The tangent modulus is estimated by laying on the curve less than 10% strain at different strain points, hence, tangent modulus data would not be clinically helpful. However, it shows at each strain point how hard or soft tissue responses are. This is relative to the micro-architecture of the tissue matrix, porosity and fluid flow rate within the matrix

It has been highlighted that compression tests are essential with modeling of viscoelasticity responses according to required tests of the United States Food and Drug Administration (FDA) and International Cartilage Repair Society (ICRS) [149,150]. Moreover, the

[148].

results in the tangent modulus of the tissue. Tangent modulus quantifies softening and hardening of the material and plastic deformation beyond yield stress [132]. Softened materials endure a higher load before ultimate failure compared to hardened materials and are suitable for replacing tissues that undergo large deformations [146]. Two factors that considerably affect the tangent modulus are strain rate and strain point. Healthy knee cartilage typically experiences average strains under 10% [36] and a maximum of 17% [147]. The tangent modulus is estimated by laying on the curve less than 10% strain at different strain points, hence, tangent modulus data would not be clinically helpful. However, it shows at each strain point how hard or soft tissue responses are. This is relative to the micro-architecture of the tissue matrix, porosity and fluid flow rate within the matrix [148].

It has been highlighted that compression tests are essential with modeling of viscoelasticity responses according to required tests of the United States Food and Drug Administration (FDA) and International Cartilage Repair Society (ICRS) [149,150]. Moreover, the American Society for Testing and Materials (ASTM) standard is focused on confined creep testing as a requirement for mechanical evaluation of designed tissues [151]. Alternatively, creep or stress relaxation is needed to quantify material properties recommended by ASTM. A systematic review of literature from 2009 to 2018 [152] showed only 11.4% of studies had performed stress relaxation or creep tests, which demonstrates that most studies did not meet the requirements of the FDA and ICRS guidance documents.

#### *7.5. Tribological Testing of Articular Cartilage and Hydrogels*

In the tribology testing of both native and engineered cartilage, there are two methods of testing the lubrication properties; the first method is sliding mate with a specified stroke length, which yields to matrix deformation. The CoF would be very low as the fluid resistance is against imposed load in the active deformation region. It is reported that the load support can be analyzed by Peclet number, where low friction occurs by the condition of Pe > > 1 and connective fluid velocity surpasses diffusive fluid velocity [5,75]. The second method of lubrication analysis is aimed more at boundary lubrication which is a stationary contact area. In this method, a sample is compressed to a solid mate, and CoF is recorded as the fluid pressure drops to the ambient pressure [81]. Therefore, interstitial fluid pressure lessens, and only contact pressure between two solid mates determines the CoF associated with the biochemical and articular surface. Thus, this method is suitable to analyze boundary lubrication and its biomolecular interactions. It is worth mentioning that a correct interpretation of using the two methods is necessary and depends on the surface and pressuring mechanism. If a tissue provides excellent permeability, which increases the localization of lubricants, it will have a relatively low CoF in stationary and high CoF in migrating contact areas. In contrast, a tissue with a remarkable pressurizing fluid mechanism but poor in boundary lubricants would have a relatively low CoF in migrating contact area and high CoF in stationary contact area [152].

#### **8. Mechanical Properties of Hydrogels**

Crosslinking process within polymer chains improves the compressive strength, stretchability and toughness of the hydrogels to withstand shear or compressive stresses [141]. There are two conventional crosslinking approaches. Covalent crosslinking enhances materials' strength and dissipates mechanical energy against deformation, whereas ionically crosslinked augments self-healing properties and controls degradation of the polymeric network [26]. Furthermore, it was reported that ionically crosslinked hydrogels using Fe3+ or Al3+ also exhibited enhanced mechanical strength [153]. Crosslinking density affects the polymer chain length, and consequently, different properties can be achieved [154]. The dangling chains phenomenon exploits the hydrophobicity and hydrophilicity interaction to form a low-crosslinked density that improves lubricious fluid retention. The high-crosslinked density, however, results in a bulk layer that enhances structure loadbearing [153]. Furthermore, interpenetrating polymer networks (IPNs) are formed by interpenetrating entanglement of two or more crosslinked polymers. A semi-IPN results

when only one polymer in the system is crosslinked, whereas, crosslinking of all polymers in the system results in full-IPN. Hence, the mechanical strengths of hydrogels in the form of full-IPN structure is superior compared to semi-IPN structures [155].

An improvement in the mechanical properties mitigates the lubrication properties of hydrogels. Subsequently, research on having a load-bearing structure with a sufficient lubricational threshold has not yielded the desired success; therefore, this subject warrants further research attention. It has been proven that monomers molar ratio, initiator and crosslinking degree determine the mechanical properties of hydrogels [156]. Zhang et al. [119] reported that the mechanical properties of bilayer hydrogels improved notably by meticulously increasing monomer (acrylic acid, AAc) content. Increasing the amount of AAc, resulted in ultimate tensile strength and elastic modulus increase. However, when AAc was more than 50%, hydrogels become very brittle and stiff, resulting in inferior tensile properties [132], and were not suitable for practical applications. Xu et al. [157] found that the titanium nanocomposite hydrogels having 10% AAc had significant tensile strength and enhanced water stability (low swelling ratio) compared to the higher molar percentage of AAc. Optimum AAc amount improves the mechanical strength and affects the nonlinearity of the hydrogels, which is a premium consideration in tissue engineering applications [158]. Arjmandi et al. [26] reported that their hydrogel's mechanical properties improved by increasing crosslinking concentration up to 21% and 32% for elastic modulus and hardness, respectively. Trivalent cations (Al3+ or Fe3+) also presented a momentous factor in increasing the strength and stiffness when hydrogels were synthesized using alginate monomer [159].

Among polymers, alginate and polyacrylamide (Alg/PAAm) have been reported to provide a high level of toughness and stretch ratio [160]. The elastic properties, furthermore, were reported to be similar to that of AC. Alg/PAAm also proved a 3-fold decrease in CoF compared to either Alg or PAAm as single network hydrogels [161]. Alg, however, has some disadvantages such as low tensile properties and difficulty in sterilization and controlling the hygiene process during synthesis. Its impurities may also affect material properties [162]. To sum up, optimum amounts of AAc, AAm, Alg and relevant crosslinking ratios would significantly improve both the mechanical and tribological properties of hydrogels.

#### *Viscoelastic and Poroelastic Relaxation*

Viscoelastic and poroelastic are associated with the rate of fluid migration within the networks, and their interaction with polymer chains results in dissipating energy [163]. Therefore, the assessment of hydrogel materials and their viscoelastic or poroelastic relaxation response is essential in designing tissues where they are subjected to high-impact loads. Hydrogels are formed by fiber networks similar to fibrin and collagen in AC and can be categorized as viscoelastic due to the exhibition of stress relaxation [164]. A nanoporous hydrogel structure, such as acrylamide hydrogels, performs minor viscoelasticity and is nearly elastic [46]. It was reported that stress relaxes promptly when the hydrogel is crosslinked ionically compared with covalently crosslinked [165]. More details were reported by Zhao et al. [165] and showed that binding and unbinding of alginate hydrogels that are crosslinked ionically show quicker stress relaxation compared to the covalently crosslinked. By exerting a force that results in unbinding of ionically crosslinked fibers, divalent cations detach from the anions of alginate chains and re-bond with another anion. In contrast, the covalently crosslinked network does not detach and re-attach fibers. Thus, instead of detaching, it yields to a longer time to relax the stress [166]. The covalently crosslinked hydrogels exhibited time-dependent mechanical properties.

It is highlighted that an abundant amount of water in hydrogels also affects viscoelastic responses. Fluid motion within the network would significantly impact dissipating energy from external loadings [165]. Hong et al. [167] formulated a coupled mass transport theory and large deformation within the hydrogel network. The motion of fluid inside the network and the resistance of the porous structure against the fluid migration yield to macroscopic mechanical relaxation, which is different from relaxation resulting from structural deformation in the network. This phenomenon is called poroelasticity and is characterized by diffusion coefficient D of the fluid in the network [168] and can be obtained by the following equation:

$$\text{D-Er2/}\mathfrak{n}\tag{1}$$

where E is the elastic modulus, r is the pore radius of the polymer network and η is the fluid viscosity in the hydrogel. According to the equation, the rate of relaxation depends on poroelasticity. As mentioned above, regarding the fluid migration, the smaller pore size results in slower fluid migration and thus slower stress relaxation. Therefore, diffusion rate D, and geometric scale L of the sample are inversely proportional to the time of stress relaxation. A smaller L yields to a faster stress relaxation due to the fluid migration at a shorter distance. However, the rate of deformations of a hydrogel is independent of the geometric scale [165]. In addition, viscoelastic responses are always attributed to fluid flow and network deformation. Therefore, when L >> √ (Dτ\_v) which was obtained for hydrogels when the sample scale is large enough to prevent the fluid from migrating to the end, viscoelastic relaxation occurs before poroelastic relaxation [169]. If we consider two states of time required for hydrogel to reach viscoelastic and poroelastic relaxation, therefore, t~τ\_v is the time of viscoelastic relaxation from deformation and t~τ\_p is the time of poroelastic relaxation resulting from fluid flow. τ\_v is the time of viscoelastic relaxation and τ\_p is the time of poroelastic relaxation. Therefore, it is essential in the design of artificial cartilage to assess the viscoelastic and poroelastic time of relaxation based on material properties.

#### **9. Tribological Properties of Hydrogels**

Wear is the loss of material, a continuous damage process due to the sliding of contact mates throughout cycles. Wear, V, is defined as the total volume of material loss. Wear rate (w) reported by Archard et al. [170] is defined as volume loss per unit sliding distance. Archard's equation predicts that the wear rate is proportional to the normal contact pressure and inversely proportional to the hardness of the material surface:

$$\mathbf{w} = \mathbf{V}/\mathbf{s} = \mathbf{K} \,\mathrm{P}/\mathrm{H} \tag{2}$$

where V is the total volume loss in [mm<sup>3</sup> ], P is the normal load in [N], H is the hardness of the material, s is the sliding distance and K is the so-called wear coefficient, a constant that is usually determined by experiment for two specific contact partners under certain environmental conditions.

A conventional system for analyzing tribological parameters is the pin-on-disk tribometer, where a small pin slides on a larger circular disk. The sliding motion is between the specimen and the rotating disk. Several types of motions and sliding between solids have been introduced (i.e., sliding wear, rolling wear, impact wear and oscillation wear) [171]. The dominant wear mechanisms are abrasion, adhesion, surface fatigue and tribochemical reactions. Abrasive wear is the subtraction of a soft material by a hard adjacent surface [172]. The most substantial part of the abrasive wear is caused by tangential sliding motions and removal of the microscopic asperities. Adhesive wear is associated with an increase in the CoF, µ between the interfaces [172]. Up to µ = 1.0, the presence of friction can be explained by adhesion itself, which means that frictional resistance is caused by asperities coming into contact and adhering to one another. Corrosive wear is a mechanism of materials and environment interface; development of worn surface may yield to different scenarios as relative motions of the bodies. Finally, wear due to fracture is a description of removal of chunks of material due to microcrack occurs within material either due to surface cracks or subsurface cracks [57].

Bilayer hydrogels that consist of a bulk layer for bearing load, and a thin porous layer to retain fluid and minimize the CoF have been developed recently [119,153]. In these

bilayer hydrogels, the bulk layer exhibited significant compressive strength up to 0.35 MPa. The reciprocating sliding test reported a 0.038 CoF associated with its lubricious layer. However, the lubricious layer was worn after a few thousand cycles due to its low network density. Surface network density is inversely proportional to water retention, which in turn influences the CoF reported by Zhang et al. [119]. Crosslinking density is proportionally related to the mesh size and showed a remarkable correlation at the transition of low to high frictions [173].

In an earlier study, the lateral and normal friction forces were not directly correlated to the stiffness but varied with the hydrogel architecture and composition [174]. The contact pressure and pore pressurization within interconnected channels are the key factors that control hydration levels in tribological assessments [168]. The contact pressure experienced by AC was reported in the range of 0.1–2.0 MPa in the hip and knee joints [76,175]. By increasing contact stress on AC, the CoF decreases [78]. However, research showed that experimental parameters and rehydration would change the trend of decreasing CoF by increasing contact stress [176].

Beyond CoF values, the determination of lubrication mechanisms in hydrogel has rarely been addressed. The effects of load and speed on lubrication regimes have been studied with the aid of the classical engineering Stribeck curve [177]. They found that hydrogels are not covered the engineering Stribeck curve regimes, and the main regimes were developed: mesh-confined, elastoviscous transition and fluid film.

In the engineering system, a prompt transition occurs over narrow ranges of the Hersey number, which is also the dimensionless fluid thickness. It is worth mentioning that stiff engineering materials have elastic moduli in the scale of GPa; therefore, the hydrodynamic fluid film would form by increasing speed or decreasing load. However, hydrogels with conformational surfaces with respect to contacting mate and a much lower range of elastic modulus (kPa) do not fall into this lubrication regime. Therefore, hydrogels are viscoelastic materials, and their wear behavior is similar to that of rubbers; thus, fatigue and adhesion wear mechanisms are dominant [178].

Furthermore, effects of applied load and sliding speed on shifting wear mechanisms have been investigated recently, and it was shown that unlike applied load, sliding speed has a minor influence on the wear mechanism [179]. Addressing these tribological properties is essential to ensure hydrogels under various contact pressures and sliding speeds can perform similar to AC [177].

A knee joint represents a situation of soft elastohydrodynamic lubrication (EHL). Artificial implants are examples of hard EHL. Hard EHL can be very successful in tribological situations, but only when the lubricating fluid has superior high-pressure rheology. This is not the case for synovial fluid, [180]; thus, the soft EHL results in thicker lubricant films than hard EHL in vivo. To this end, a porous architecture of polymer would mimic natural cartilage in terms of EHL lubrication and yields significant performance to conventional fully dense polymers. The CoF associated with different polymeric materials are presented in Table 4.


**Table 4.** Effects of monomers and polymers materials on hydrogels' CoF.


**Table 4.** *Cont.*

\*Stc: Static; \*Dyn: Dynamics; \*PAMPS: Poly 2-acrylamido−2-methyl−1-propanesulfonic acid; \*PVA-c: Poly vinyl-alcohol cryogel; \*PAMPS: Poly (2-acrylamide−2-metyl-propane sulfonic acid) and polyacrylamide; \*PAMPS/PDMAAm: Poly-(2-Acrylamido−2-methylpropane sulfonic acid)/poly-(N,N'-dimetyl acrylamide).

#### **10. Strengthening of Hydrogels with Nanoparticles**

TiO<sup>2</sup> nanoparticles (TiO<sup>2</sup> NPs), due to their low toxicity, excellent biocompatibility, low cost and high-level stability, have been explored for the synthesis of polymeric hydrogels for medical applications [185]. However, due to the hydrophobic nature of these nanoparticles, having a homogenous solution that affects mechanical and tribological properties would be challenging. These challenges are because of TiO<sup>2</sup> NPs surface and electrostatic attraction among particle molecules [186]. TiO<sup>2</sup> nanoparticles tend to agglomeration or aggregation due to solution ionic strength (IS), pH level, surface charge or coating [187]. Using different techniques may affect the tendency of nanoparticles to aggregation. Some researchers have reported these techniques, which are ultrasonic irradiation, stabilize TiO<sup>2</sup> NPs in an aqueous medium, electrostatic stabilization, controlling pH level of the solution by neutralizing acidity level, and coating the surface of nanoparticles by surfactants [188]. Moreover, overcoming the van der Waals attraction of nanoparticles by utilizing steric or electrostatic stabilization is the critical factor to suppress nanoparticle aggregation or agglomeration effectively.

Ultrasonic irradiation was an effective method to disperse NPs, which depends on solvent type, concentration and suspension volume. Two ultrasonic irradiation methods, bath and probe sonications, are commonly used, although probe sonication showed a better result [187]. Even using probe sonication is not the permanent solution to suppress aggregation. Stabilizers were reported to have prolonged effects on dispersed particles [189]. As mentioned earlier, steric and electrostatic stabilization takes place when charges accumulate by the particle surfaces. More than 30 mV or less than −30 mV surface charge on the TiO<sup>2</sup>

NPs yields no aggregation. Moreover, having higher than 1% TiO<sup>2</sup> NPs concentration in the AAm-based hydrogels composition resulted in sedimented particles even if a long homogenization process was used [190].

Some monomers of hydrogel compositions have a high acidity level, for instance, AAc, which can affect NPs dispersion. The hydrodynamic size of nanoparticles can be tuned by modifying the pH level of the solution. TiO<sup>2</sup> and SiO<sup>2</sup> particles have a positive surface charge when the pH level is low, and on the opposite, negative surface charge when the pH level is high [191].

## *TiO<sup>2</sup> and Silica NPs Mechanical and Tribological Properties*

Chemically crosslinked co-polymeric hydrogels have been reported to have superior mechanical properties compared to the conventional chemically crosslinked homopolymeric hydrogels. This has been attributed to the formation of more uniform and compact networks in the co-polymeric hydrogels [154]. However, nanocomposite hydrogels, loaded with nano particles (NCHs), reported higher strength, improved sliding wear resistance, anisotropy and potential self-healing property compared with double-network hydrogels (DNHs), topology hydrogels (TPHs) and micromolecular microsphere hydrogels (MMHs). The swelling ratio is a crucial factor for hydrogels in biomedical applications, which supports water-stability within the hydrogel and can be achieved by utilizing titania NPs [154]. The superior mechanical strength of hydrogels is associated with the equilibrium swelling state. Seddiki et al. [133] reported that TiO<sup>2</sup> NPs and a high dosage of crosslinking agents (15%) are vital factors affecting swelling ratio. Furthermore, it has been reported that carboxyl groups formed complexes with TiO<sup>2</sup> NPs via different methods to crosslink polymer chains [192].

The concentration of TiO<sup>2</sup> NPs is a critical point in the reinforcement process since this substrate act as a crosslinker. The higher concentration of NPs, which is inversely proportional to the structure mesh size, would produce a higher degree of crosslinking [190]. Consequently, with smaller mesh-size, hydrogels would imbibe less fluid in the networks, which affects stress distribution over the structure. Due to this fact, the poroelasticity and viscoelasticity relaxation time would also be affected.

Silica nanoparticles (SNPs) have also been utilized to synthesize artificial cartilage and have demonstrated appreciable mechanical and biological properties [143,193]. Incorporating SNPs within polymer networks improves tissue adhesion, stiffness and shear modulus [194]. Furthermore, SNPs, interlaced with polymer chains, enhances hydrogel elasticity [195]. Zareie et al. [196] showed that by increasing SNPs amounts in the polyacrylamide networks, the number of tie points in each entanglement increased, and the compressive strength of hydrogel reached 26.2 kPa.

In addition to improving mechanical strength, SNPs have promoted the degree of crosslinking in very weak chemically crosslinked PAAm hydrogels, which have interestingly presented the ability of SNPs to function as a crosslinker [195]. Arjmandi and Ramezani [146] reported that SNPs interact with PAAm chains resulting in network crosslinks through hydrogen bonds.

Unlike other NPs, SNPs showed a significant impact on initial shear modulus and viscoelastic properties since they could immobilize the polymer chains and form NPspolymer interphases [197]. SNPs reported increasing the number of tie points in each entanglement, which results in the improvement of the compressive strength [196]. SNPs also enhance slower chain kinetics and relaxation due to tough NPs-polymer bonds [146]. Polymer bonds relax promptly when NPs are located far from chains [198]. Viscoelasticity of the SNP loaded nanocomposite hydrogels (NCHs) was studied extensively and found to be similar to that of AC [195]. AC exhibits a time-dependent response associated with viscoelasticity, poroelasticity or the combination of both phenomena [43,199].

Tribologically, SNPs showed the dominance of adhesion mechanisms rather than other wear mechanisms, although fatigue wear took place with surface pitting at higher applied loads [200]. Utilizing 1–4% SNPs into the PAAm-alginate network resulted in low Cof values in the range 0.0035–0.0055, which is comparable to the CoF of AC (0.0001) [146]. It is attributed to the strong interfacial NPs-polymer bonding in the hydrogel matrix. The contact pressure and pore pressurization within interconnected channels are the key factors that control hydration levels in tribological assessments [168]. SNPs also affect mesh patterns, and therefore, are strongly correlated with the lubrication regimes [200].

## **11. Conclusions**

In this paper, a comprehensive review of the literature for the AC is presented. First, the architecture of the AC, its compositions and the role of each component on mechanical and tribological properties were discussed extensively. It was explained that damaged cartilage cannot recover itself due to its avascular nature. Then, osteoarthritis roots and treatment methods were presented with conventional TKR/THR solutions as the ultimate treatment being highly invasive and with significant disadvantages especially for younger patients, and the need for revision surgery due to the limited service life of TKR/THR implants were discussed. To address the gap in treatment of younger patients with OA, developments of artificial cartilage by different synthesizing processes, materials and their pros and cons were described. The required standard and necessary tests for artificial cartilage to assess its mechanical and tribological properties based on the International Cartilage Repair Society (ICRS), Food and Drug Administration (FDA) and American Society for Testing and Materials (ASTM) were briefly reviewed. Viscoelastic properties were found as the critical point in the design of engineered soft tissues and the techniques to tune viscoelasticity to perform optimum responses under different loading scenarios were reviewed. Advanced bilayer hydrogels were discussed as a promising candidate for artificial cartilage. Both the load-bearing and lubricious layer were investigated recently; however, the weak point of the proposed lubricious layer was found to be its limited strength and service life under cyclic sliding tests.

Polymeric hydrogels have indeed provided a practical alternative to AC when OA treatment is considered. This is largely attributed to the progress attained in finding the appropriate combinations of materials as well as techniques for the synthesis of hydrogels with mechanical and biochemical properties of natural cartilage. Polymeric hydrogels stand to emerge as an attractive technology for AC replacement applications. Even though highly promising, the application of hydrogels in AC replacement are not free from challenges of biocompatibility. It is, therefore, imperative that attention be diverted to understanding the degradability of synthetic polymeric networks and the interaction of the hydrogels with cells in physiological conditions. A further area of innovation would be addressing the fabrication challenges of hydrogels which will make them safer and ready for clinical use. The mechanical properties of hydrogels are an important consideration for AC replacement application; however, mechanical characterization of hydrogels have been limited to mainly cell free scaffolds. Cells' encapsulation can potentially dictate profound changes in mechanical properties of hydrogels. Hence, mechanical characterization of cell seeded hydrogels should be a consideration for the future.

**Author Contributions:** Conceptualization, M.M.; writing—original draft preparation, M.M.; writing —review and editing, A.N. and M.R.; supervision, A.N. and M.R. All authors have read and agreed to the published version of the manuscript.

**Funding:** This research received no external funding.

**Institutional Review Board Statement:** Not applicable.

**Informed Consent Statement:** Not applicable.

**Data Availability Statement:** The data presented in this study are available on request from the corresponding author.

**Conflicts of Interest:** The authors declare no conflict of interest.

## **References**

