Smart Orthopedic Biomaterials and Implants

Abstract

Musculoskeletal injuries including bone defects continue to present a significant challenge in orthopedic surgery due to suboptimal healing. Bone reconstruction strategies focused on the use of biological grafts and bone graft substitutes in the form of biomaterials-based 3D structures in fracture repair. Recent advances in biomaterials science and engineering have resulted in the creation of intricate 3D bone-mimicking structures that are mechanically stable, biodegradable, and bioactive to support bone regeneration. Current efforts are focused on improving the biomaterial and implant physicochemical properties to promote interactions with the host tissue and osteogenesis. The “smart” biomaterials and their 3D structures are designed to actively interact with stem/progenitor cells and the extracellular matrix (ECM) to influence the local environment towards osteogenesis and de novo tissue formation. This article will summarize such smart biomaterials and the methodologies to apply either internal or external stimuli to control the tissue healing microenvironment. A particular emphasis is also made on the use of smart biomaterials and strategies to create functional bioactive implants for bone defect repair and regeneration.

Graphical Abstract

Stimuli used in smart biomaterial design to induce bone repair/regeneration. Smart biomaterials can enhance the development and growth of new bone tissue due to their response to various stimuli. These biomaterials are organized based on the location of their stimuli in terms of the defect site (internal and external). Some stimuli can be delivered both internally and externally, which gives rise to a novel category of internally and externally responsive smart biomaterials.

1. Introduction

Bone fractures—a global public health issue––originate from trauma, fractures, autoimmune and aging diseases, tumor resection, and surgical complications [1]. It has been reported that in the United States there are approximately 15 million bone fractures each year, which includes 1.6 million traumatic fracture hospital admissions, 2 million osteoporotic fractures, 500,000 knee replacements, and 350,000 hip replacements [2, 3]. Altogether, bone defects have a substantial cost of over 60 billion dollars to the healthcare industry and require an estimate of 1.6 million bone grafts annually [2, 3, 4]. These numbers are expected to rise due to an increase in life expectancy and the baby boomer population [4]. Therefore, there is a tremendous need to develop novel biomaterials to enhance bone fracture repair and regeneration treatment options [5, 6].

The traditional and commonly used materials for bone defect treatment include autografts, allografts, xenografts, decalcified bone matrices, bioceramics, and metals [7]. However, the use of these materials in Orthopedic surgery is limited by problems concerning their clinical availability, regenerative potential, and further medical issues. Autografts are deemed to be the “gold standard” for bone regeneration but have supply limitations and the possibility to cause significant donor site morbidity, inflammation, pain, and resorption of the implanted bone [2, 8]. Allografts and xenografts are common substitutes for autografts; however, they face immunogenicity concerns, donor shortages, and disease transmission [9, 10]. Metallic implants have the potential to corrode and release cytotoxic ions that can cause inflammation and allergic responses leading to tissue loss [11]. To reduce the possible foreign-body response with metallic implants, researchers have coupled bioceramic materials such as varying forms of calcium phosphates including hydroxyapatite with these implants; however, bioceramics bring additional problems due to their brittleness and resorption rates [9, 11]. Ultimately, these current methods of clinically treating bone defects can match the mechanical and structural properties of bone to provide a mechanically stable environment for initial tissue development and growth but fail to be biomimetic, which prevents the regeneration and formation of functionalized tissue [12].

The design and development of biodegradable polymers as biomaterials led to the creation of a tissue engineering paradigm [13, 14, 15]. Even though early efforts were limited to biodegradable polymer development, later on the focus was to design strategies for tissue repair and regeneration. In this direction, numerous natural and synthetic biodegradable biomaterials alone or in combination with stem/progenitor cells and osteogenic growth factors were utilized to develop bone tissue engineering strategies [16, 17, 18]. Researchers have fabricated 3-dimensional (3D) scaffolds with interconnected pore structures for cell infiltration and new tissue formation. Our group designed and developed scaffold systems from biodegradable polymers, polymer blends, and composites with ceramics, carbon nanotubes, and biodegradable metals [19, 20, 21]. These optimally porous and biomechanically compatible scaffold systems have shown excellent cell compatibility and growth in vitro and tissue compatibility and vascularized bone tissue formation in vivo [5, 22, 23, 24].

To further introduce functionality into biomaterials, researchers have focused on the use of smart biomaterials. Smart/intelligent biomaterials can be defined as a biomaterial that could positively influence tissue regeneration through physical/chemical or electrical/magnetic stimulation. The idea is to either externally or internally stimulate the progenitor/stem cells via the biomaterial implant to induce new bone tissue formation, also known as osteogenesis [25]. The incorporation of these stimuli has been achieved utilizing advanced biomaterials and/or fabrication methods [25]. In this review, recent smart biomaterial trends and designs will be highlighted, along with the various stimuli used to form smart bone biomaterials/grafts. The review also provides a new classification for smart biomaterials and presents insights into the future of smart biomaterials and their use in designing better implants for bone defect repair and regeneration.

2. Internal Stimuli

As discussed earlier, current clinically used biomaterials fail to create an environment that accurately represents the bone tissue where the biomaterial is implanted, so it is unable to regenerate functionalized tissue. An approach that counters this issue involves functionalizing biomaterials to respond to the internal microenvironment surrounding the defect site [26, 27]. The internal microenvironment includes various chemical and biological features that can be applied to the biomaterial’s composition to enable stimuli-responsive bone regeneration. Chemical stimuli found in the microenvironment include pH, surface chemistry interactions, and specific molecules like glucose, while biological cues are enzymes and bioactive molecules. Another main characteristic of the internal microenvironment is its temperature, which can be utilized in stimulating shape-memory biomaterials. These specific internal stimuli are illustrated and summarized in Figure 1. Some of the well-studied examples for each stimulus are highlighted in Table 1.

Figure 1:

Schematic of different types of internal stimuli used in smart biomaterials and their mechanisms. (a) An acidic pH is associated with the microenvironment of a bone defect. pH-responsive biomaterials contain acidic functional groups, such as COOH, which can alter their morphology and dimensions in response to decreases in local pH. (b) Enzyme responsive smart biomaterials incorporate substrates of enzymes whose activity and expression are linked to pathologic microenvironments. Upon implantation, interactions between substrates and target enzymes cause changes in the biomaterial structure, leading to release of loaded drugs or growth factors. (c) Shape-memory biomaterials undergo deformation in response to changes in temperature. Before implantation, these biomaterials are subject to cooling, which allows for deformation to improve ease of implantation and initial interactions with the defect site. After implantation, the internal body temperature brings the implant to its transition temperature, where it will return to its original shape.

Table 1.

Summary of Internal Stimuli-Responsive Biomaterials including Chemical, Biological, and Shape-Memory stimuli examples.

Stimuli	Biomaterial/Scaffold (Fabrication Method)	Effect of Stimuli	Ref.
Chemical Stimuli	pH-triggered, multilayered osteogenic and antimicrobial coating (3D-Printing)	Ion release for osteogenesis and antimicrobial activity in vivo	[*31]
	GelMA-oxidized sodium alginate hydrogels loaded with gentamicin sulfate (GS) and phenamil (Phe) (UV Crosslinking)	pH responsiveness allowed for controlled release of GS and Phe in vivo	[32]
	Bisphosphonate-based nanocomposite hyaluronic acid hydrogel (HA-BP) (UV Crosslinking)	Release of BP in response to acidic pH - Regulates osteoclast activity for enhanced bone regeneration in vivo	[33]
	Biomimetic micro/nano-scale (MNS) titania fiber-like network (Alkaline Treatment)	Induced osteogenic differentiation of progenitor cells and M2 polarization of macrophages in vivo	[36]
	Glucose-sensitive controlled-release fiber scaffold (Electrospinning)	Structural changes in response to elevated levels of blood glucose - Controlled release of rhBMP-2 in vivo	[*38]
	GL13K antimicrobial peptide and an MMP-9 responsive peptide coating (Co-immobilization)	The co-immobilized peptide surface showed potent anti-biofilm activity - effective osteoblast and fibroblast proliferation in vitro	[*41]
	Polydopamine-modified Ti substrates with surface nanoparticles (Layer-by-Layer Assembly)	The enzymatic response showed antimicrobial and osteogenic activity in vitro	[42]
Biological Stimuli	Deferoxamine (DFO)-loaded titania nanotubes (TNT) substrates (Layer-by-Layer Assembly)	Antimicrobial activity and promotion of progenitor cell proliferation/differentiation in vitro	[43]
	ROS responsive hydrogel coating loaded with thymosin β4 (Tβ4) (Anodic Oxidation)	ROS responsive degradation and release of Tβ4 -promoted osteogenic activity in vitro and in vivo	[*47]
	Gelatin-chitosan multilayer encapsulating SEW2871 (Layer-by-Layer Assembly)	Controlled release of SEW2871 - Improved osteointegration in vivo	[51]
Shape-Memory Biomaterial	Nitinol (Ni-Ti) arched shape-memory connector (ASC)	Provided continuous concentration compression to accelerate osseous healing and restoration in a clinical study	[57]
	Ni-Ti shape memory alloy scaphoid arc nail fixator (NT-SAN)	Generated continual fixation under body temperature - improved healing in a clinical study	[58]
	Porous poly (ε-caprolactone)-diols (PCL-diols) scaffold coated with hydroxyapatite (Particle Leaching)	Enhanced cell attachment and osteogenesis in vivo	[60]
	3D porous PEG/PTMG/citrate functionalized amorphous calcium phosphate (CCACP) (Sequential Gas Foaming and Freeze-Drying)	Minimally invasive in vivo implantation and accelerated formation of bone-like apatite	[61]
	Aniline trimer (AT)-based, self-healable conductive polyurethane scaffold (Oxygen Plasma Surface Modification)	Demonstrated self-healing in vivo	[*62]

2.1. Chemical Stimuli

Chemical signals within the defect microenvironment can be utilized to initiate biomaterial response to regenerate functionalized bone tissue. While there are various types of chemical stimuli that are present in smart biomaterial design, this review focuses on microenvironment pH, surface chemistry, and specific molecules that trigger the action, such as glucose.

Typically, the pH of extracellular organelles and the bloodstream is 7.4; however diseased, inflamed, or infected tissues often present a more acidic microenvironment [28]. This change in microenvironment pH has been targeted by the pH-responsive biomaterials that alter their morphology and dimensions to release therapeutic and/or regenerative compounds [26, 29]. Acidic pH-responsive biomaterials are derived from polymers that contain acidic functional groups, such as -COOH and -SO₃H [30]. Some of the commonly used pH-responsive polymers include hyaluronic acid, chitosan, alginic acid, gelatin poly(acrylic acid), and poly(l-glutamic acid) [30]. For instance, Deng et al. [*31] fabricated porous polyetheretherketone (PEEK) scaffolds with a coating that provided the biomaterial with a novel bacteria-triggered, acidic pH-responsive property. The coating is described as a unique polydopamine (pDA) and silver nanoparticle (AgNP) sandwich (pDA-Ag-pDA), where AgNPs were trapped within the first pDA layer and apatite is anchored onto the second layer. When the scaffold interacts with an acidic microenvironment provided by a bacteria-contaminated bone defect, Ag⁺ ions are released immediately to eliminate the bacteria in the defect, while Ca²⁺ and PO₃⁴⁻ are released rapidly to provide osteogenic cues to the surrounding progenitor cells. This study is a proof of concept that a biomaterial coating can induce osteogenesis — or be osteoinductive — and therapeutic at the same time.

Similarly, Yao et al. developed gelatin methacryloyl (GelMa) – oxidized sodium alginate (OSA) hydrogels for pH-responsive dual controlled release of an antibiotic, gentamicin sulfate (GS), and a BMP-2 activator, phenamil (Phe) [32]. The study utilized an activator of BMP-2 rather than the protein itself to reduce the dosage of exogenous BMPs, which can have serious side effects for the patient. Phe was encapsulated into mesoporous silicate nanoparticles (MSN) to achieve delayed and controlled release of the BMP-2 activator in response to decreasing pH. Controlling the release kinetics of loaded therapeutic and osteoinductive molecules is a desired characteristic of smart bone biomaterials, which is significant for future designs involving large bone defect models. MSNs are known to have excellent pH sensitivity, biodegradability, and high drug loading efficiency. The hydrogel showcased efficient antibacterial activity, induced osteogenic differentiation of C2C12 cells, and significantly promoted new bone formation in a criticalsized cranial bone defect.

Other current research on the topic has looked toward creating biomaterials that respond to the acidity induced through osteoclast activity [33]. This research by Li et al. utilizes a pH-responsive, bisphosphonate-based nanocomposite hyaluronic acid (HA-BP) hydrogel that responds to the acidic microenvironment created by osteoclasts during the negative feedback mechanism of bone homeostasis with the triggered release of bisphosphonates (BP). BPs have inhibitory effects on osteoclast activity and bone resorption signaling, which can be used to emulate the negative feedback properties of native bone and homeostasis, ultimately enhancing bone regeneration. The HA-BP hydrogel allowed for preosteoclastic differentiation and activation of macrophages, which is needed for bone remodeling, but inhibited osteoclastic maturation through the triggered release of BP during the production of acidic osteoclast factors. The HA-BP hydrogel displayed increased bone regeneration in vivo as well. The success of the experiment provides new pathways for pH-responsive biomaterials for bone regeneration as it demonstrates a design that is used in a greater population of patients dealing with bone defects since it targets the natural process of bone homeostasis. Traditionally, studies focus on the acidic pH in the defect microenvironment; however, this study highlights how the acidity of standard osteoclast function can be used as a stimulus for its own inhibition, which would encourage consistent bone formation.

The chemical composition of the biomaterial, specifically its surface, is a notable characteristic that controls the adsorption of molecules in the microenvironment and tunes cellular adhesion, growth, and differentiation [34]. During fabrication, the surface chemistry of the biomaterial can be modified by incorporating various types of ions and/or bioactive compounds that can improve biological responsiveness and promote tissue regeneration [34, 35]. In a study by Bai et al., titanium implants were fabricated to have a biomimetic coating derived from a micro/nano-scale (MNS) titania fiber-like network [36]. The MNS coating created a negative surface charge on the titanium, which when placed in alkali solution results in the creation of an apatite layer which increases bioactivity. The new surface chemistry of the titanium induced BMSC osteogenic differentiation and polarization of M1 macrophages to the pro-healing M2 phenotype. The implant surface chemistry showed significant enhancement of de novo bone formation and integration with the existing bone, or osseointegration. By establishing a bioactive surface for a biomaterial, interactions and adsorption of molecules within the microenvironment can lead to increased integration of the implant, as well as enhanced bone regeneration.

Besides pH and surface chemistry, researchers have investigated the use of specific molecules, such as glucose, as stimuli for bone regenerative biomaterials based on either their presence or absence [35]. Glucose has been specifically analyzed for this purpose due to the link between patients with diabetes and alveolar bone loss [37]. Glucose-responsive biomaterials typically contain glucose oxidase, phenylboronic acid, or lectin within the structure as they can interact with fluctuating levels of glucose to deform the biomaterial to release therapeutic and/or regenerative components [37]. Jiang et al. [*38] created a novel glucose-sensitive controlled-release fiber scaffold that was grafted with glucose oxidase and contained recombinant human bone morphogenetic protein 2 (rhBMP-2) to promote bone regeneration. In response to fluctuating levels of glucose, the scaffold transforms glucose into gluconic acid, which in turn allows for the controlled release of rhBMP-2 since the scaffold’s physical properties are affected by the local pH change. In a diabetic rat bone defect model, it was seen that this scaffold slowly released rhBMP-2, which promoted bone tissue regeneration. The concepts explored in this experiment can be used to develop novel biomaterials that are stimulated by the signature molecules seen in the pathology of other common bone loss diseases.

2.2. Biological Stimuli

Biological cues within a bone defect microenvironment are the main contributors to implant failure and infection. However, these cues can be easily targeted and used as stimuli for smart biomaterials. Common biological cues that can be targeted using biomaterials include enzymes, increased ROS, and bioactive molecules.

Dysregulation of enzyme activity and expression is known to be linked to the pathology of various diseases including cancer, osteoarthritis, and osteoporosis [39, 40]. Due to these properties, researchers have recently attempted to target enzymes in their smart biomaterial approaches to take advantage of signaling pathways that already exist within the cells and tissues in the microenvironment [*41, 42, 43]. The most common types of enzymes targeted for smart biomaterials are proteases, lipidases, oxidoreductases, and glycosidases [35, 44]. More specifically, matrix metalloproteinases (MMPs) are often used for these purposes [44]. Fischer et al. [*41] created a titanium bone implant that can respond to MMPs and bacteria due to a co-immobilized surface containing GL13K antimicrobial peptide and an MMP-9 responsive peptide. MMP-9 was targeted during this study since it is highly expressed during osteoclast-mediated bone remodeling, as well as initial implant site clearing. MMP-9 cleavable peptides also can be combined with bone regenerative peptides, allowing for greater bone formation and healing. When introduced to the environments containing osteoblasts, osteoclasts, and bacteria, the co-immobilized surface showed simultaneous response to the enzyme and microbial activity, enabled osteoblast and fibroblast proliferation, and inhibited biofilm formation. Similar to the previously mentioned study by Li et al. [33], this study explores ways to target osteoclast activity to prevent their function at an implant site and enhance early bone formation and osteointegration.

In a comparable study, Yu et al. [43] developed enzyme-responsive titanium substrates with antimicrobial properties. These substrates were designed to target hyaluronidase, which is a glycosidase that is secreted by bacteria that are commonly found on and around the surface of bone implants causing implant failure. The titanium implant utilized deferoxamine (DFO)-loaded titania nanotube (TNT) substrates with polyelectrolyte multilayers of hyaluronic acid (HA)-gentamicin (Gen) conjugates (HA-Gen). The HA-Gen conjugates can resist bacterial adhesion and display antimicrobial properties as it is degraded by hyaluronidases. The degradation of the HA-Gen multilayers allows for the controlled localized release of DFO, which is used to encourage osteogenic and angiogenic differentiation around the implant site. During cellular and bacterial studies, the substrates demonstrated successful adhesion, proliferation, osteogenic/angiogenic differentiation of MSCs, as well as effective antibacterial properties. These studies suggest an alternative approach for bone regeneration where bacterial enzymes found within the bone defect can be used as stimuli for anti-microbial and regenerative purposes.

ROS are essential signaling molecules within the body that are produced mainly from the mitochondria during physiological processes [45, 46]. Similar to acidic pH and dysregulated enzymes, abnormal levels of ROS are also seen in various disease environments [45]. For bone-related diseases, higher states of ROS are linked with osteoporosis, rheumatoid arthritis, and bone metastases [45, 46]. The main forms of ROS that promote osteoclast activity during these diseases are superoxide (O₂) and hydrogen peroxide (H₂O₂) [46]. Biomaterials can be constructed to respond to high levels of ROS by either allowing the ROS to alter its properties, mainly solubility, or by having ROS-mediated chemical properties that induce bond cleavage reactions [45]. In a study by Li et al. [*47], ROS found within a femoral bone defect is targeted by a Tβ4-loaded titanium implant that is coated with a multifunctional hydrogel. The hydrogel coating included a borate ester bond, which can be rapidly oxidized in the presence of H₂O₂, therefore providing the implant with ROS-responsive degradation characteristics. In response to the high levels of ROS in the defect site, the implant released Tβ4, which utilized the ROS-induced inflammatory response to promote the transformation of macrophages from the M1 phenotype to M2. This alteration in phenotype inhibits fibrosis around the implant and encourages the osteogenic differentiation of MSCs. This in vivo study highlighted the implant’s osteointegration and vascularization within the defect site. Currently, immunomodulatory biomaterials have been of great interest in their applications for bone regeneration as they take advantage of normal body processes involved in the immune response. This research provided a novel method for inducing this effect compared to the common use of bioactive molecules, which is explained below.

Other common biological stimuli used in the development of smart biomaterials are bioactive molecules, such as antigens or macrophage recruiting agents. To target antigens, biomaterials are incorporated with antibodies toward the target antigen [48]. At defect sites, antigens are overexpressed and allow for targeted delivery of regenerative and/or therapeutic agents when utilizing antibodies within the smart biomaterial design [48]. Similarly, macrophages can be targeted during the immune response that can occur when a biomaterial is implanted into a defect site. Macrophages can be grouped into two different types, M1 and M2, where M1 macrophages are pro-inflammatory and M2 are pro-remodeling [49]. By developing smart biomaterials with immunomodulatory properties that can coordinate M1–M2 macrophage polarization, it is possible to take advantage of the body’s immune response to enhance osteointegration and bone tissue regeneration [50]. In a study by He et al. [51], this immunomodulatory effect was obtained by using a micro-structured titanium implant that encapsulated SEW2871 in a gelatin-chitosan multilayered coating. SEW2871 is a macrophage recruitment agent and was slowly released by the biomaterial as the multilayers degraded. Adding to this, the multilayer coating design can be attributed to the biomaterial’s ability to polarize macrophages due to the changes in wettability and difference in protein adsorption between the layers. It was observed that the biomaterial facilitated MSC osteogenic differentiation in vitro and improved osteointegration when implanted in vivo. The ability to control macrophage recruitment and polarization has been of great interest in smart biomaterials to help design biomaterials. This approach utilizes biological mechanisms involved during implantation that can cause device rejection/failure and manipulate its function to promote new bone tissue formation.

2.3. Shape-Memory Biomaterials

Shape-memory biomaterials (SMBs) can be intentionally deformed but recover their original shape in response to a stimulus [52]. SMBs are designed to respond to internal temperature within a defect site; however, other stimuli such as light, electric fields, and magnetic fields are currently being evaluated as well [53]. For this review, we will focus on the studies involving thermoresponsive SMBs. Currently, SMBs have gained much research attention due to their clinical applications in minimally invasive surgery [25, 52, 54]. Before implantation, SMBs are deformed by cooling to ease implantation and then expanded to fill the bone defects in response to heating via exposure to physiological temperature [25, 52, 54]. This shape-modifying ability is highly advantageous for clinical applications as it allows for ease of implantation with initially minimized form/shape that can eventually unfold to achieve the desired shape and functionality for the defect at hand [54]. Three main types of SMBs are developed, these include shape memory alloys (SMAs), shape memory polymers (SMPs), and composites derived from SMA and SMP [52, 55].

SMAs are typically derived from either titanium-nickel alloys (Nitinol) or nickel-free titanium alloys. Nitinol SMAs, especially when porous, possess many advantageous characteristics for bone regeneration, such as high strength, low stiffness, high toughness, good mechanical stability, and biocompatibility [52]. Thermoresponsive SMAs possess the shape memory effect (SME), the transformations between stable configurations within martensite and austenite phases due to the change in temperature [54, 55, 56]. The martensite phase is formed when the metal is quenched rapidly from the austenite phase and is known to be soft, flexible, and stable at low temperatures [55, 56]. Oppositely, austenite is stable at high temperatures and makes the metal hard and stiff [55, 56]. Xia et al. [57] studied the clinical use of nitinol (Ni-Ti) SMA within displaced olecranon fractures. The study utilized naturally arched shape-memory connectors (ASCs) to improve the treatment methods for olecranon fractures. The ASC design possesses high strength and low modulus, occupies less space causing minimal patient discomfort, and has simpler implantation/removal procedures. The device had a transition temperature of 33°C and was malleable at lower temperatures when in the martensite phase. Before implantation, the device would be cooled at 0–4°C for plastic deformation. After implantation, 40–50°C water was used to stimulate the device to employ its shape-memory function when returning to the austenite phase. After implantation, the ASC device is subjected to internal compression due to movement and gravity. While this causes problems for other devices used for olecranon fractures, the ASC device can continuously evenly distribute the compression forces to utilize them in the healing process. It was observed that out of the 57 patients with varying types of olecranon fractures, only 9 patients had postoperative complications and none of the patients requested to remove the ASC device, which is a major problem with current clinical methods. Also, the average healing time of the fracture was 15 weeks, where all patients showed an anatomical or nearly anatomical reduction of the olecranon fracture. These results showed that the device accelerated osseous healing and restoration, while decreasing the need for secondary operations and patient discomfort.

In a similar clinical study, Song et al. utilized Ni-Ti SMA as a scaphoid arc nail fixator (NT-SAN) [58]. The implant was fabricated in a way to meet the anatomical characteristics of the scaphoid and biomechanical requirements of a scaphoid fracture in terms of fixation and fatigue strengths, respectively. Matching the previously discussed study, before implantation the NT-SAN is cooled by 0–4°C to induce deformation and then reheated with 40°C saline after implantation to begin recovery through SME. The NT-SAN had notable ease of implantation and improved the healing rate for scaphoid fractures while causing less scaphoid damage compared to other clinically used methods. These clinical trials provide data on how smart biomaterials have the potential to open new doors with tissue regeneration strategies.

Recently, research on SMBs has shifted greatly towards SMPs due to their advantages over SMBs. While SMPs do not have the mechanical strength and recovery stress of SMA, they are lightweight, low-cost, easily processable, biodegradable, more corrosion resistant, possess excellent shape change and have tunable properties [52, 59]. The SME of thermoresponsive SMP consists of the fixed phase and the reversible phase, where the fixed phase maintains the original shape and the reversible phase fixes the temporarily deformed shape in response to the transition temperature (T_Trans) [53, 60]. In other words, when the temperature is above T_Trans, the polymer becomes more malleable [52, 53, 60]. Upon cooling, the shape becomes fixed, but when the temperature increases above T_Trans again, the programmed molecular chain will restore the desired shape [61]. Huang et al. [60] fabricated a porous SMP scaffold using poly (ε-caprolactone)-diols (PCL-diols), which was coated with hydroxyapatite to increase the osteogenic potential. The porous structure of the scaffold has been a notable addition to recent SMP studies as it minimizes the invasiveness of the scaffold and encourages tissue in-growth, so this design characteristic will be expected in most future research in the field. Through the combination of the porous structure and HA coating, BMSC adhesion, proliferation, spreading, and osteogenic differentiation were enhanced. It was also observed that the SMP scaffold was able to maintain its porous structure through implantation and fully degrade in 6 months in vivo while supporting tissue invasion and vascularization.

Shaabani et al. [*62] developed a self-healing SMP derived from an aniline trimer (AT)-based polyurethane scaffold (PU-AT). This study focused on utilizing an SMP to create a scaffold that can heal itself in response to damage caused by implantation to further improve SMP’s ability to be minimally invasive. When implanting the temporary, compressed version of SMP, micro-cracks can be found within the structure if the materials used are not flexible enough. This can lead to structural failure of the implant causing more problems for the patient. When the scaffolds were subjected to a razor blade scratch in vitro, self-healing of the scaffolds was observed when the scaffolds were heated at 40°C. This is due to the synergistic effect of shape memory properties and gold-thiol bonds. This innovative design shows potential in increasing the regenerative lifetime for biomaterials and this concept should be explored within other areas of smart biomaterials that are facing implantation-based deformities. The PU-AT scaffold also showed excellent osteogenic potential both in vitro and in vivo. Osteogenic differentiation of human ASCs increased, along with osteoblast proliferation and adhesion. When implanted into a calvarial defect model, the PU-AT scaffold displayed a significant increase in bone regeneration compared to the control.

3. External Stimuli

While internal stimuli mainly respond to pathological conditions, some defect conditions require a targeted delivery of therapeutic agents and/or growth factors with specific dosage and timing. Using an external source of stimuli, dosage and release times can be controlled by altering the activity and intensity of the applied stimulus. Two commonly used external stimuli for smart biomaterials include magnetic fields and light. Typically, these stimuli can be integrated into the design of a biomaterial by including magnetic nanoparticles (MNPs) and photothermal nano-agents, respectively [26]. Below, magnetically active, light-responsive smart biomaterials are described including some of the recent studies that have successfully utilized each external stimulus (Figure 2) (Table 2).

Figure 2:

Graphical overview of mechanisms for external stimuli. (a) Demonstration of how magnetic stimuli encourage bone regeneration. Magnetically active biomaterials are designed to induce regeneration in response to external magnetic fields applied to the defect site. These biomaterials incorporate iron oxide nanoparticles into the design due to their superparamagnetic behavior and ability to increase calcium content within the defect site. (b) Visualization of light as a smart biomaterial stimulus. In response to an applied controlled light source, such as UV and NIR, light-responsive biomaterials adjust their polarity, hydrophobicity, and/or confirmation. These characteristic changes for the biomaterial can modulate cell behavior and guide loaded factors to influence bone tissue regeneration.

Table 2.

Summary of External Stimuli-Responsive Biomaterials including Magnetically active and Light responsive biomaterial systems.

Stimuli	Biomaterial/Scaffold (Fabrication Method)	Effect of Stimuli	Ref.
Magnetically Active	Silk fibroin and poly(2-hydroxyethyl methacrylate) scaffold template with Magnetite (Fe3O4) nanoparticles (Free Radical Polymerization)	Increased cell proliferation and osteogenic differentiation in vitro	[63]
	Titanium substrates with CoFe2O4/P(VDF-TrFE) nanocomposite coatings (Tape Casting)	Improved osteogenic differentiation, adhesion, and proliferation of preosteoblasts in vitro	[64]
	Macroporous gelatin sponges loaded with SDF-1a and alginate ferrogels loaded with BMP-2 (Cryopolymerization)	Ability to control the release of BMP-2 in vitro by selecting the time at which a magnetic stimulus is applied	[*65]
	Nanohydroxyapatite-based composite co-doped with iron oxide (IO) nanoparticles and functionalized with miR-21 and miR-124 (Microwave-Stimulated Hydrothermal Method)	Allowed for controlled release of functional mRNAs in vitro	[66]
	Magnetic Iron oxide (α-Fe2O3 and γ-Fe2O3) nanoparticles (Sol-Gel)	Enhanced osteogenic differentiation in vitro, modulation of integrins	[67]
Light Responsive Biomaterial	Graphene oxide-loaded chitosan hydrogel film containing Teriparatide (Electrodeposition)	NIR-activated Teriparatide release, induced local bone regeneration in vivo	[71]
	Poly(N-isopropylacrylamide-co-nitrobenzyl methacrylate) (pNIPAm-co-NBMA) microgels loaded with dexamethasone (DEX) (Free Radical Precipitation Polymerization)	UV light controlled osteogenic differentiation in vitro	[*72]
	Titanium implant coated with TiO2 nanoparticles doped with fluorine/dopamine/collagen (Hydrothermal Method)	NIR stimulation can simultaneously ablate osteosarcoma and promote the osteogenic activity of BMSCs in vitro	[*74]
	In situ generated calcium phosphate nanoparticle (ICPN)-coordinated poly(dimethylaminoethyl methacrylate-co-2-hydroxyethyl methacrylate) (DHCP) hydrogel loaded with PTH (Sol-Gel and Water-in-Oil Emulsion)	NIR stimulation triggered an on-demand release of PTH, as well as micropore formation in situ	[75]
	Fibrin hydrogel and gold nanoparticles loaded with genetically modified cells with NIR responsive, rapamycin-dependent gene switch (Galvanic Replacement)	BMP-2 production was induced by cell constructs in vitro that converted energy from the NIR laser into heat	[76]

3.1. Magnetically Active Biomaterials

Magnetically active biomaterials utilize external magnetic fields or direct magnetic forces to encourage bone tissue regeneration [63, 64, *65, 66, 67]. The current trend for magnetically responsive biomaterials is to incorporate iron oxide nanoparticles into the design due to their magnetic properties and bone regeneration applications [37, 68]. Iron oxide nanoparticles that have a size of <100 nm showcase superparamagnetic behavior, which is beneficial for the magnetic responsiveness of the biomaterial and prevents particle agglomeration [69]. These nanoparticles will randomly flip their orientation of magnetic moments in response to thermal fluctuations, which prevents the retention of magnetism after the applied magnetic field is removed [63, 69]. It is hypothesized that iron oxide nanoparticles within the scaffolds can induce bone regeneration by increasing the calcium content in the defect site when an external magnetic field is applied [63].

Tanasa et al. [63] embedded magnetite nanoparticles with this effect into the silk fibroin scaffolds to induce and accelerate the osteogenic process in response. When a 120 mT magnetic field was applied, it was observed that the embedded magnetic nanoparticles influenced the orientation of actin filaments and the distribution of cells around the scaffold. Also, stimulation from an external magnetic field increased the cellular proliferation potential and positively impacted the osteogenic potential of preosteoblasts. A significant takeaway from the experiments conducted in this study is that they demonstrated that neither the magnetic field nor scaffold alone contributed to the osteogenic effect, but a smart synergistic system composed of both factors encouraged these results.

Tang et al. [64] took a different approach and developed a magnetic responsive bioactive coating that can be added to titanium substrates used for bone implants. In this approach, the magnetically active coating was derived from CoFe₂O₄ nanoparticles that were combined into a P(VDF-TrFE) matrix. The study compared the activity of preosteoblasts that were seeded onto titanium substrates with different concentrations of CoFe₂O₄ within the coating. It was found that while all the coatings significantly increased the adhesion, proliferation, and osteogenic differentiation of preosteoblasts, the low magnetization coating (6% CoFe₂O₄) had the highest ALP activity and gene expression. This experiment showcased that the nanoparticles form of iron oxide is not a requirement for magnetically active biomaterials, which provides the potential for more design concepts that utilize an iron oxide coating.

As mentioned earlier, an important property of biomaterials using external stimulation is that the dosage and release times of loaded factors can be controlled by changing the activity of the stimulus. In a study by Madani et al. [*65], this property is showcased using a two-compartment cylindrical hydrogel system that can respond to magnetic stimulation. The outer compartment of the system consisted of macroporous gelatin sponges (gelfoam) loaded with SDF-1a, a recruitment factor of MSCs, while the inner compartment consisted of magnetically responsive alginate ferrogels loaded with BMP-2, an osteogenic factor. The study aimed to find the optimal dose and release timing of BMP-2 that correspond to the release of SDF-1a. The two-compartment design of the biomaterial allows for the rapid release of SDF-1a, to recruit MSCs towards the hydrogel, and a controlled release of BMP-2 based on applied magnetic stimulation. The research demonstrated that controlled release of BMP-2 was possible with the application of an external magnetic stimulus, which increases the clinical applications for magnetically active biomaterials. The delayed release of BMP-2 after SDF-1a initiated MSC recruitment and migration towards the biomaterial system significantly enhanced the degree of osteogenic differentiation.

3.2. Light Responsive Biomaterials

Light is another stimuli source that has been targeted in smart biomaterials for bone regeneration. Current research has utilized various types of light including ultraviolet (UV), visible, near-infrared (NIR), and lasers of differing wavelengths [54]. Biomaterials that are commonly used to respond to these light sources include carbon-based nanomaterials (such as graphene oxide and carbon nanotubes), gold-based nanomaterials, graphite carbon nitride, transition metal oxides (mainly TiO₂), transition metal dichalcogenides, nanosheets, and nanoparticles [70]. Light-responsive biomaterials are advantageous since the intensity, wavelength, duration, and location of the light source can be precisely controlled in a non-invasive manner [28, 69, 70, 71]. In response to light, these biomaterials will adjust their polarity, hydrophobicity, and confirmation to modulate cell behavior, guide loaded nano-vehicles, and regulate the defect area microenvironments [30, 70].

UV light is typically used in light-responsive biomaterial systems that are designed to induce drug release from the biomaterial due to the short wavelength of UV light [69, *72]. In a study by Zhang et al. [*72], UV light was used to facilitate the release of dexamethasone (DEX), a synthetic glucocorticoid osteogenic inducer, from a microgel to induce osteogenic differentiation of hMSCs. Long-term systemic injections of DEX are correlated with a higher risk of osteoporosis and osteonecrosis, so the efforts are to achieve controlled and precise release of the drug through UV light stimulation. In this study, the microgel was synthesized with O-nitrobenzyl ester, which is the most commonly used photocleavable molecule and is known for membrane destabilization and drug release upon light irradiation [13]. It was observed that the DEX-loaded microgels induced hMSC osteogenic differentiation during UV light exposure. The progress of differentiation can be controlled by turning the light source on and off, which provides a function that can easily be translated clinically.

NIR light irradiation of light-responsive biomaterials has become a focus of researchers because of UV light’s limited tissue penetration and harmful effects on biological tissue [28, 69, 71, 73, *74, 75, 76]. Adding to this, NIR light is known to be able to deeply penetrate tissues due to its minimum refraction and reflection by biomolecules and chromophores [73]. Wang et al. [71] utilized NIR light to reduce the graphene oxide (GO) that was loaded into a chitosan hydrogel film containing Teriparatide, to take advantage of the osteogenic potential and drug delivery properties of GO. GO is a photothermally active material that can serve as a drug reservoir without decreasing the loaded drug’s potency and efficacy. In this study, GO was used as a carrier for Teriparatide that can release the drug through a photothermal heating effect when triggered by NIR irradiation. A hand-held NIR light emitter was used by the researchers throughout the study to control the duration and delivery rate of NIR to determine the most effective pulsatile delivery schedule. It was shown that pulsatile release of Teriparatide in vivo can induce local bone regeneration and angiogenesis within an osteoporotic bone defect compared to a constant release, but further work is needed to determine the most effective pulsatile schedule. The use of a pulsatile schedule for delivery shows the potential to incorporate automation within the design process, which could enhance the regenerative effect.

NIR responsiveness was also employed in a biomaterial engineered by Wu et al. [*74] to combat common problems in bone defect repair, such as incomplete removal of tumor cells and insufficient osseointegration. The group fabricated a titanium implant that was coated with TiO₂ nanoparticles doped with fluorine, dopamine (PDA), and collagen. TiO₂ nanoparticles can be used for photothermal therapy (PTT) and photodynamic therapy (PDT) when NIR light is applied. Both therapies can be applied to tumor treatment and have a synergistic effect on cancer cell ablation when combined. PTT utilizes thermal effects to rupture cell membranes and denature proteins/enzymes of tumor cells, whereas PDT produces ROS that restricts the progression of tumor invasion. By doping the TiO₂ nanoparticles with PDA and collagen, the applications of the biomaterial can also include the promotion of osteogenic differentiation within the defect site. PDA has outstanding photothermal properties and can graft bioactive molecules due to its abundant groups. In response to NIR light, PDA and collagen are modified which grants the Ti implant excellent osteogenic activity. In vitro experiments using Saos-2 and BMSCs displayed the synergistic effect of PTT and PDT that can eliminate osteosarcoma cells within 10 minutes, while the PDA and collagen on the implant surface enhance the proliferation and differentiation of BMSCs. This study demonstrated the potential to link regenerative biomaterials with therapeutic treatments. By creating a multifunctional smart biomaterial with these characteristics, patient recovery can progress more rapidly as tissue regeneration can begin during disease treatment.

4. Internally and Externally Responsive

Currently, smart biomaterials and their stimuli have been categorized in many ways, but the most common method is to base smart biomaterials on the location of the applied stimuli, internal or external, for the implant (Graphical Abstract). However, in some cases, specific types of stimuli are found both internally or externally depending on the biomaterial used and the aim of the design. Thus, the overlap between internally and externally responsive biomaterials lead to a new category of “internally and externally responsive” biomaterials. Piezoelectric biomaterials fall in this category as they respond to internally experienced mechanical stresses and externally applied ultrasonic or electromagnetic fields [77, 78, 79, 80]. In this section, piezoelectric biomaterials/implants are discussed in terms of this new classification with notable examples (Figure 3) (Table 3).

Figure 3:

Graphical demonstration of the processes involved in piezoelectric smart biomaterials. (a) Once implanted, piezoelectric biomaterials can be stimulated by either internal mechanical forces or an exterior source, such as LIPUS, that influences the direct piezoelectric effect. (b) In response to the stimulus, the implant generates a local electric field due to the direct piezoelectric effect. (c) The electrical cues generated by the implant is transferred to cells and tissue within the defect, which can stimulate cell signaling pathways. (d) The promotion of cell signaling pathways leads to enhanced growth factor synthesis that enhances cell proliferation, osteogenic differentiation, and osteogenesis.

Table 3.

Summary of Piezoelectric Biomaterials that utilize direct or converse piezoelectric effect

Stimuli	Biomaterial/Scaffold (Fabrication Method)	Effect of Stimuli	Ref.
Piezoelectric	ZnO nanoparticle/PVDF composite fiber membranes (Electrospinning)	Osteoblast proliferation in vitro enhanced due to piezoelectric effect from the scaffold	[*84]
	Porous PVDF scaffolds coated with hydroxylated BaTiO3 nanoparticles functionalized with polydopamine (Selective Laser Sintering)	Electric cues significantly promoted cell adhesion, proliferation, and differentiation in vitro	[85]
	Polycaprolactone-tricalcium phosphate (PCL-TCP) films coated with PVDF (Solvent-Casting)	Osteoinduction in vitro by both PVDF and PEMF	[86]
	Porous titanium/titanium alloy scaffolds coated with BaTiO3 (Electron Beam Melting)	Promoted osteogenic differentiation of BMSCs in vitro, improved the osteoinductive activity of implants in vivo	[*87]
	Whitlockite (Ca18Mg2(HPO4)2(PO4)12) nanoparticles (WH NPs) (Co-Precipitation)	WH NPs produce electric fields that promote osteoblast proliferation and differentiation in vitro when stimulated by LIPUS.	[88]

4.1. Piezoelectric Biomaterials

Bone is known to have natural piezoelectric properties, which are used for bone remodeling and repair [77, 78]. Researchers have manufactured piezoelectric biomaterials that can mimic the natural piezoelectric behavior of bones to enhance the bone regeneration process [77, 79, 80]. Piezoelectric biomaterials can either generate an electric charge in response to an applied mechanical stress (direct piezoelectric effect) or can provide mechanical stimulation in response to electrical stimuli (converse piezoelectric effect) [78, 79, 80]. Smart biomaterials that utilize the direct piezoelectric effect have great potential in bone regeneration since the electrical stimulation that is generated by the biomaterial is transferred to surrounding cells and tissue, which promotes cell signaling pathways leading to enhanced growth factor synthesis [79]. Piezoelectric biomaterials are constructed from piezoceramics (Barium titanate [BaTiO₃], zinc oxide [ZnO], piezoelectric polymers (polyvinylidene fluoride [PVDF], PVDF-TrFE, and poly-l-lactic acid [PLLA]), or polymer/ceramic or glass/ceramic composites [78, 79, 81, 82]. BaTiO₃ is the most studied piezoceramic since it is highly biocompatible and has high piezoelectric coefficients [63, 64]. While PVDF is the most widely used piezoelectric polymer for bone tissue engineering due to its excellent mechanical properties, biocompatibility, and processability [83].

Internally responsive piezoelectric biomaterials are designed to restore the electrophysiological microenvironment of bone upon implantation into the defect site [77]. These biomaterials respond to the applied mechanical loads and generate an electric field, which can serve as stimuli for bone regeneration. Xi et al. [*84] electrospun ZnO/PVDF composite fibers to fabricate an internally responsive piezoelectric scaffold. By incorporating ZnO into the electrospun fibers, the β-phase crystal content of PVDF was further increased, ultimately improving the piezoelectric properties of the composite scaffold. When the piezoelectric scaffold was tested under mechanical strain, human osteoblast proliferation was significantly promoted by the local electric field that was generated. The scaffold also was able to inhibit the growth of bacteria in vitro due to attraction to the cationic surface of the scaffold and interactions with high concentrations of ZnO NPs.

A related study by Shuai et al. [85] utilized a piezoelectric composite derived from a PVDF matrix and BaTiO₃ NPs. Similar to ZnO in the previous example, BaTiO₃ significantly increased the fraction of β-phase PVDF within the scaffold’s microstructure. The BaTiO₃ NPs were functionalized with polydopamine, which allowed for a uniform distribution. The interactions between polydopamine and BaTiO₃ NPs attributed to an increase in β-phase PVDF enhanced the output voltage by 356%, which promoted adhesion, proliferation, and differentiation of human osteosarcoma fibroblasts. Additionally, the tensile strength and modulus of the piezoelectric scaffolds increased, furthering their potential for bone regeneration usage. These studies demonstrate how tissue regenerative techniques can incorporate a patient’s movements to establish a more practical treatment. By applying force to the defect site, which can be done in daily activities or monitored movements, patients can experience regenerative effects due to stimulation of the piezoelectric biomaterial.

For some bone defects and diseases, patients are debilitated and do not provide the natural mechanical or electrical stimulation needed for piezoelectric biomaterials to effectively promote bone regeneration [86]. These situations require piezoelectric biomaterials to be stimulated from an exterior source that induces mechanical forces or electrical stimuli [86, *87, 88]. A study by Liu et al. [*87] utilized low-intensity pulsed ultrasound (LIPUS) to externally induce the piezoelectric effect. LIPUS stimulated a porous Ti alloy scaffold coated with BaTiO₃ (BTi) to increase the osteoinductive activity both in vitro and in vivo. LIPUS can be used as an external mechanical force to induce BaTiO₃ to create an electric field, as well as to further promote bone regeneration. The BTi scaffolds were compared against porous Ti alloy (pTi) scaffolds which are known to maintain mechanical stability within a defect site and provide space for tissue ingrowth within the porous structure but are not bioactive leading to long-term failure of the scaffold.

It was hypothesized that the BTi scaffolds would promote bone regeneration by restoring the physiological electrical microenvironment of the defect site in response to mechanical stimulation. This was proved through in vivo micro-CT and Van-Gieson stain bone volume analysis (BV/TV) of both pTi and BTi scaffolds (Figure 4). The micro-CT images showed that after 8 months of implantation, bone tissue (yellow) had almost filled the pores of each scaffold; however, the BTi scaffold was significantly more integrated into the surrounding bone (red) than the pTi scaffold (Figure 4a). It was also observed that de novo bone formation significantly increased inside the BTi scaffold than in the pTi scaffold at both 4 and 8 months, which can be seen qualitatively when comparing the scaffold’s BV/TV (Figure 4c). Van-Gieson staining displayed the osteoinductive effect of each scaffold by highlighting continuous and intact de novo bone formation (red) (Figure 4b). Again, BV/TV was used to quantitatively express bone volume, and at 4 and 8 months the BTi group had a significantly higher BV/TV than pTi (Figure 4d). The study demonstrated that the direct piezoelectric effect generated from BaTiO₃ through external stimulation promoted osteogenic differentiation of BMSCs in vitro and de novo bone formation within a sheep femoral defect model, which provides great potential for developing effective fracture repair/regeneration methods.

Figure 4:

Porous titanium alloy scaffold coated with BaTiO₃ to develop an externally responsive piezoelectric biomaterial. (a) 2D (left) and 3D (right) micro-CT analysis of pTi and BTi scaffolds after 8 months of in vivo implantation. (b) Van-Gieson stained histological sections of pTi and BTi after 8 months of in vivo implantation. (c-d) BV/TV analysis of pTi and BTi scaffolds after 4 and 8 moths of in vivo implantation, asterisks (*) indicate significance. Adapted from the study by Liu et al [70] with permission of SAGE Publications, copyright 2020.

Dong et al. [86] used pulsed electromagnetic fields (PEMF) to stimulate piezoelectric polycaprolactone-tricalcium phosphate (PCL-TCP) films coated with PVDF. PEMF can be used to induce the converse piezoelectric effect by electrically stimulating piezoelectric scaffolds to deliver mechanical stimuli within the defect site. Adding to this, PEMF can also stimulate progenitor or MSC cell population in the defect area. Therefore, by combining PEMF’s regenerative abilities with the piezoelectric properties of scaffolds, it is possible to create an effective alternative strategy for bone tissue repair.

Experimentally, it was observed that PEMF and the piezoelectric scaffold have a synergistic effect on the proliferation and differentiation of preosteoblasts in vitro. This design differs from commonly studied piezoelectric biomaterials as it is based off the converse piezoelectric effect, which opens up new clinical treatment modalities by utilizing smart piezoelectric implants.

5. Conclusion and Future Directions

Bone grafting has evolved to meet the needs of reconstruction ranging from nondegradable bone cement to biodegradable biomaterials-based bone graft substitutes with superior clinical utility. Research on the cell-material interaction in response to topography, stiffness, porosity, and functional groups laid the foundation for modulating cell behavior. Integrating these features into biomaterials and their 3D structures results in smart/intelligent biomaterials and implants. Changes to physiological parameters including pH, chemical milieu, and temperature can act as internal stimuli and activate biological, chemical, and shape memory effects to smart biomaterials. Alternatively, smart bone implants can be activated using an external source such as light or magnetic field stimuli to promote osteogenesis. The piezoelectric biomaterial systems can be activated by both external and internal stimuli by the application of mechanical load or LIPUS to produce charge separation. It is evident from literature reports that both internal/external stimulation using smart biomaterial systems promote cell-material interaction and vascularized bone formation in animal models.

Significant progress has been made in the creation and application of smart biomaterials for bone repair with improved healing outcomes. However, underlying mechanisms that cause improved cell behavior and osteogenesis are poorly understood. For example, the mechanisms for smart biomaterial LIPUS-induced charge separation and improved osteoblast performance are unclear. Another challenge is to determine internal/external stimuli strength and treatment duration required for a specific bone defect repair and regeneration. As such, these studies are still in their infancy and more data in preclinical models are needed for a clinical translation. An active area of research is to apply smart biomaterials in the repair and regeneration of other musculoskeletal tissues such as cartilage, tendon, muscle, and ligament. Cellular mechanisms that promote tissue regeneration under various stimuli will revolutionize the future of smart engineered implants and tissue regeneration.